Biointegration of Medical Implant Materials: Science and Design

Biointegration of medical implant materials

© Woodhead Publishing Limited, 2010

Related titles: Cellular response to biomaterials (ISBN 978-1-84569-358-9) The response of cells to biomaterials is critical in medical devices. It has been realised that specific cell responses may be beneficial – encouraging adhesion, healing or cell multiplication. Cellular response to biomaterials will examine the response of cells with a wide range of materials, targeted at specific medical applications. The first section reviews cellular response to polymers and ceramics. A second group of chapters discuss cell responses and regenerative medicine for nerves, muscles and orthopaedic materials. The final chapters analyse the effect of surface chemistry and how it can be manipulated to provoke a useful cell response. Surfaces and interfaces for biomaterials (ISBN 978-1-85573-930-7) This book presents the current level of understanding on the nature of a biomaterial surface, the adaptive response of the biomatrix to that surface, techniques used to modify biocompatibility, and state-of-the-art characterisation techniques to follow the interfacial events at that surface. Tissue engineering using ceramics and polymers (ISBN 978-1-84569-176-9) Tissue engineering is rapidly developing as a technique for the repair and regeneration of diseased tissue in the body. This authoritative and wide-ranging book reviews how ceramic and polymeric biomaterials are being used in tissue engineering. The first part of the book reviews the nature of ceramics and polymers as biomaterials together with techniques for using them such as building tissue scaffolds, transplantation techniques, surface modification and ways of combining tissue engineering with drug delivery and biosensor systems. The second part of the book discusses the regeneration of particular types of tissue from bone, cardiac and intervertebral disc tissue to skin, liver, kidney and lung tissue. Details of these and other Woodhead Publishing materials books can be obtained by: • •

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© Woodhead Publishing Limited, 2010

Biointegration of medical implant materials Science and design Edited by Chandra P. Sharma

Oxford

Cambridge

© Woodhead Publishing Limited, 2010

New Delhi

Published by Woodhead Publishing Limited, Abington Hall, Granta Park, Great Abington, Cambridge CB21 6AH, UK www.woodheadpublishing.com Woodhead Publishing India Private Limited, G-2, Vardaan House, 7/28 Ansari Road, Daryaganj, New Delhi – 110002, India www.woodheadpublishingindia.com Published in North America by CRC Press LLC, 6000 Broken Sound Parkway, NW, Suite 300, Boca Raton, FL 33487, USA First published 2010, Woodhead Publishing Limited and CRC Press LLC © Woodhead Publishing Limited, 2010 The authors have asserted their moral rights. This book contains information obtained from authentic and highly regarded sources. Reprinted material is quoted with permission, and sources are indicated. Reasonable efforts have been made to publish reliable data and information, but the authors and the publishers cannot assume responsibility for the validity of all materials. Neither the authors nor the publishers, nor anyone else associated with this publication, shall be liable for any loss, damage or liability directly or indirectly caused or alleged to be caused by this book. Neither this book nor any part may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, microfilming and recording, or by any information storage or retrieval system, without permission in writing from Woodhead Publishing Limited. The consent of Woodhead Publishing Limited does not extend to copying for general distribution, for promotion, for creating new works, or for resale. Specific permission must be obtained in writing from Woodhead Publishing Limited for such copying. Trademark notice: Product or corporate names may be trademarks or registered trademarks, and are used only for identification and explanation, without intent to infringe. British Library Cataloguing in Publication Data A catalogue record for this book is available from the British Library. Library of Congress Cataloging in Publication Data A catalog record for this book is available from the Library of Congress. Woodhead Publishing ISBN 978-1-84569-509-5 (book) Woodhead Publishing ISBN 978-1-84569-980-2 (e-book) CRC Press ISBN 978-1-4398-3064-2 CRC Press order number: N10181 The publishers’ policy is to use permanent paper from mills that operate a sustainable forestry policy, and which has been manufactured from pulp which is processed using acid-free and elemental chlorine-free practices. Furthermore, the publishers ensure that the text paper and cover board used have met acceptable environmental accreditation standards. Typeset by Toppan Best-set Premedia Limited, Hong Kong Printed by TJ International Limited, Padstow, Cornwall, UK

© Woodhead Publishing Limited, 2010

Contents

Contributor contact details Preface

xi xv

1

Biointegration: an introduction C. K. S. Pillai and C. P. Sharma, Sree Chitra Tirunal Institute for Medical Sciences and Technology, India

1

1.1 1.2 1.3 1.4 1.5

Introduction Biointegration of biomaterials for orthopedics Biointegration of biomaterials for dental applications AlphaCor artificial corneal experience Biointegration and functionality of tissue engineering devices Percutaneous devices Future trends References

1 1 7 8

1.6 1.7 1.8

Part I Soft tissue biointegration 2

2.1 2.2 2.3 2.4 2.5 2.6 2.7 2.8

Biocompatibility of engineered soft tissue created by stem cells P. A. Clark, University of Wisconsin–Madison, USA; and J. J. Mao, Columbia University, USA Introduction Bone: from tissue to molecular organization Bone development Bone homeostasis Bone repair after injury Bone and joint disease Current treatment options and total joint replacements Current challenges of titanium implants

10 10 11 12

17

19

19 20 22 24 26 29 29 30 v

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vi

Contents

2.9 2.10

Current titanium modifications for improved integration Mimicking nature toward achieving titanium ‘biointegration’: cytokines and implants Growth factor delivery: why is controlled and sustained release important? Future trends Acknowledgements Sources of further information and advice References

2.11 2.12 2.13 2.14 2.15 3

3.1 3.2 3.3 3.4

Replacement materials for facial reconstruction at the soft tissue–bone interface E. Wentrup-Byrne, Queensland University of Technology, Australia; L. Grøndahl and A. Chandler-Temple, The University of Queensland, Australia

32 34 35 37 39 39 40

51

Introduction Facial reconstruction Materials used in traditional interfacial repair Surface modification of facial membranes for optimal biointegration Future trends Acknowledgements References

51 55 60

4

Corneal tissue engineering Y.-X. Huang, Ji Nan University, China

86

4.1 4.2 4.3

Introduction Characteristics of the human cornea and its regeneration Special conditions for wound healing and tissue regeneration of the cornea Approaches to corneal tissue engineering Future trends References

86 87

3.5 3.6 3.7

4.4 4.5 4.6

73 78 78 79

90 95 108 109

5

Tissue engineering for small-diameter vascular grafts J. I. Rotmans, Leiden University Medical Centre, The Netherlands; and J. H. Campbell, University of Queensland, Australia

116

5.1 5.2

Introduction Required characteristics of tissue engineered blood vessels

116 118

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vii

5.3 5.4 5.5 5.6

Approaches to vascular tissue engineering Future trends Conclusion References

120 136 138 138

6

Stem cells for organ regeneration K. D. Deb, Dayananda Sagar Institutions, India

147

6.1 6.2 6.3

Introduction Basic components of tissue engineering Tissue engineering and stem cells in organ regeneration Conclusions References

147 149

6.4 6.5

Part II Drug delivery 7

7.1 7.2 7.3 7.4

Materials facilitating protein drug delivery and vascularisation P. Martens, A. Nilasaroya and L. A. Poole-Warren, University of New South Wales, Australia

161 169 169

177

179

Introduction Hydrogel classification Factors influencing protein encapsulation and release Tissue engineering applications: vascularisation and protein delivery Conclusions Acknowledgements References

179 181 185

8

Inorganic nanoparticles for targeted drug delivery W. Paul and C. P. Sharma, Sree Chitra Tirunal Institute for Medical Sciences and Technology, India

204

8.1 8.2 8.3 8.4 8.5 8.6 8.7

Introduction Calcium phosphate nanoparticles Gold nanoparticles Iron oxide nanoparticles Conclusion Acknowledgements References

204 206 213 218 226 226 227

7.5 7.6 7.7

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191 197 198 198

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Contents

9

Alginate-based drug delivery devices L. Grøndahl, G. Lawrie and A. Jejurikar, The University of Queensland, Australia

236

9.1 9.2 9.3 9.4 9.5 9.6

Introduction Alginate biopolymers Drug delivery using alginate matrices Future trends Acknowledgement References

236 237 247 258 259 259

10

Functionalised nanoparticles for targeted drug delivery S. Manju and K. Sreenivasan, Sree Chitra Tirunal Institute for Medical Sciences and Technology, India

10.1 10.2 10.3 10.4 10.5 10.6

Introduction Drug targeting Multifunctional nanocarrier systems Conclusion Acknowledgements References

Part III Design considerations 11

11.1 11.2 11.3 11.4

11.5 11.6 12

12.1 12.2

Biocompatibility of materials and its relevance to drug delivery and tissue engineering T. Chandy, 3M Drug Delivery Systems, USA Biocompatibility of materials and medical applications Biomaterials for controlled drug delivery Biomaterials for tissue engineering Role of the scaffold and loaded drug/growth factor in the integration of extracellular matrix and cells at the interface Future trends References Mechanisms of failure of medical implants during long-term use A. Kashi and S. Saha, SUNY Downstate Medical Center, USA Introduction Manufacturing deficiencies

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267

267 269 282 289 289 290

299

301 301 307 311

315 320 321

326

326 327

Contents

ix

12.3 12.4 12.5 12.6 12.7 12.8 12.9 12.10

Mechanical factors (e.g. fatigue, overloading) Wear Corrosion Clinical factors for implant success and failure Failure mechanisms of non-load-bearing implants Failure analysis of medical implants Conclusion References

328 329 333 334 335 337 340 342

13

Rapid prototyping in biomedical engineering: structural intricacies of biological materials S. J. Kalita, University of North Dakota, USA

349

13.1 13.2 13.3 13.4 13.5 13.6 13.7 13.8 13.9 13.10 13.11

Introduction An overview of biomaterials Material properties of structural biomaterials Rapid prototyping – a novel manufacturing approach Designing structural implants Rapid prototyping in biomedical engineering – synopsis Rapid prototyping in mimicking structural intricacies of biological materials Patient-specific customized scaffolds via rapid prototyping Conclusion List of abbreviations References

377 387 387 389 390

Index

399

© Woodhead Publishing Limited, 2010

349 354 357 360 369 373

Contributor contact details

(* = main contact)

Chapter 2

Editor C. P. Sharma Division of Biosurface Technology Biomedical Technology Wing Sree Chitra Tirunal Institute for Medical Sciences and Technology Poojappura Thiruvananthapuram 695 012 India Email: [email protected]

Chapter 1 C. K. S. Pillai and C. P. Sharma* Division of Biosurface Technology Biomedical Technology Wing Sree Chitra Tirunal Institute for Medical Sciences and Technology Poojappura Thiruvananthapuram 695 012 India Email: [email protected]

P. A. Clark University of Wisconsin–Madison UW-Hospitals and Clinics Department of Neurological Surgery CSC K4/879 600 Highland Ave. Madison, WI 53792 USA Email: [email protected] J. J. Mao* Columbia University College of Dental Medicine Fu Foundation School of Engineering and Applied Sciences Department of Biomedical Engineering 630 W. 168 St. – PH7 East CDM New York, NY 10032 USA Email: [email protected]

xi © Woodhead Publishing Limited, 2010

xii

Contributor contact details

Chapter 3 E. Wentrup-Byrne Visiting Fellow Tissue Repair and Regeneration Program Institute of Health and Biomedical Innovation Queensland University of Technology 60 Musk Avenue, Kelvin Grove Brisbane, QLD 4059 Australia Email: [email protected]

Chapter 4 Y.-X. Huang Institute of Biomedical Engineering Ji Nan University Guang Zhou China 510632 Email: [email protected]

J. H. Campbell Australian Institute for Bioengineering and Nanotechnology Corner College and Cooper Rds (Bldg 75) The University of Queensland Brisbane, QLD 4072 Australia Email: [email protected]

Chapter 6 K. D. Deb Tissue Engineering and Regenerative Medicine Lab Department of Biotechnology Dr C. D. Sagar Center for Life Sciences Dayananda Sagar Institutions Shavige Malleswara Hills Kümarawamy Layout Bangalore 560078 India Email: [email protected], [email protected]

Chapter 5 J. I. Rotmans* Leiden University Medical Centre Department of Nephrology, C3-P Albinusdreef 2, 2333 ZA Leiden The Netherlands Email: [email protected]

Chapter 7 P. Martens*, A. Nilasaroya and L. A. Poole-Warren Graduate School of Biomedical Engineering University of New South Wales Sydney, NSW 2052 Australia Email: [email protected], [email protected]

© Woodhead Publishing Limited, 2010

Contributor contact details

xiii

Chapter 8

Chapter 11

W. Paul and C. P. Sharma* Division of Biosurface Technology Biomedical Technology Wing Sree Chitra Tirunal Institute for Medical Sciences and Technology Poojappura Thiruvananthapuram 695 012 India

T. Chandy 3M Drug Delivery Systems Division 3M Center Building 260-03-A-06 St. Paul, MN 551344-1000 USA

Email: [email protected]

Chapter 12

Chapter 9

A. Kashi* SUNY Downstate Medical Center 450 Clarkson Avenue, Box 30 Brooklyn, NY 11203 USA

L. Grøndahl*, G. Lawrie and A. Jejurikar School of Chemistry and Molecular Biosciences The University of Queensland Cooper Rd, Brisbane, QLD 4072 Australia Email: [email protected], [email protected]

Chapter 10 S. Manju and K. Sreenivasan* Laboratory for Polymer Analysis Biomedical Technology Wing Sree Chitra Tirunal Institute for Medical Sciences and Technology Poojappura Thiruvananthaparam 695 012 India Email: [email protected]

Email: [email protected]

Email: [email protected] S. Saha Department of Orthopaedic Surgery and Rehabilitation Medicine SUNY Downstate Medical Center 450 Clarkson Avenue, Box 30 Brooklyn, NY 11203 USA Email: [email protected]

Chapter 13 S. J. Kalita Engineered Surfaces Center School of Engineering and Mines University of North Dakota 4201 James Ray Drive, Suite 1100; Stop 8391 Grand Forks, ND 58202 USA Email: [email protected]. edu

© Woodhead Publishing Limited, 2010

Preface

The aim of this book is to enhance our understanding of the interfacial interaction and integration of implant materials with hard/soft tissue. The ultimate success of any implant inside the body is not only that it should be non-toxic and biocompatible with respect to its physico-chemical properties, including degraded products if any, for a desired application, but also that it integrates with the tissue as per the physiological requirements. Such issues have been discussed appropriately by experts in three separate sections on soft tissue biointegration, drug delivery and design considerations. The book will certainly be useful for academic faculty graduate students and the medical devices industry interested in understanding the concepts useful for enhancing the quality of their products. I thank all the authors who contributed the chapters in this book and Ms Lucy Cornwell for coordinating the communication link among all of us. I would also like to thank our former Director Prof. K. Mohandas, Dr K. Radhakrishnan Director SCTIMST and Dr G. S. Bhuvaneshwar Head BMT Wing SCTIMST Trivandrum for providing facilities to complete this project. Chandra P. Sharma

xv © Woodhead Publishing Limited, 2010

1 Biointegration: an introduction C. K. S. P I L L A I and C. P. S H A R M A, Sree Chitra Tirunal Institute for Medical Sciences and Technology, India

Abstract: This introductory chapter discusses the relevance of the basic aim of the book, reflecting on the importance of the interfacial concepts related to the biointegration of the implants with respect to their design and structure–function relationship. Various examples, including orthopedic implants, are discussed. Key words: biointegration, interface, medical implants, design, biomaterials.

1.1

Introduction

When biomaterials are designed, a set of properties are built in such a way as to ensure that, after implantation, they will help the body to heal itself. So it is of critical importance that these materials be integrated into organspecific repair mechanisms such as the physiologic process required for the biologization of implants (Amling et al., 2006). It should involve a direct structural and functionally stable connection between the living part and the surface of an implant. Although various materials have been developed in recent years with enhanced physical, surface and mechanical properties, the use of these materials in certain biological applications is often limited by poor tissue integration. So, the question is on how biomaterials can be converted to ‘living tissues’ after implantation. To cite an example, the bonding of hydroxyapatite to bone, which is considered as a true case of biointegration, is thought to involve a direct biochemical bond of the bone to the surface of an implant at the electron microscopic level and is independent of any mechanical interlocking mechanism (Meffert et al., 1987; Cochran, 1996). Several groups working on various aspects of the design, development and application of improved devices are concerned with how these materials become integrated with soft and hard tissues in the body and how these implanted systems have to match their physical–chemical and biological properties to those of their environment.

1.2

Biointegration of biomaterials for orthopedics

Biomaterials are defined as ‘materials intended to interface with biological systems to evaluate, treat, augment or replace any tissue, organ, or function 1 © Woodhead Publishing Limited, 2010

2

Biointegration of medical implant materials

of the body’ (Williams, 1999). Orthopedic research is developing and advancing at a rapid pace as new techniques are applied to musculoskeletal tissues. The discovery of biologic solutions to important problems, such as fracture-healing, soft-tissue repair, osteoporosis, and osteoarthritis, continues to be an important research focus. At the same time, research on biomaterials and biomechanics is critical for advances in current areas such as tissue-engineering and cytokine delivery. In orthopedic applications, there is a significant need and demand for the development of a bone substitute that is bioactive and exhibits material properties (mechanical and surface) comparable with those of natural, healthy bone. Particularly, in bone tissue engineering, nanometer-sized ceramics, polymers, metals and composites have been receiving much attention recently. This is a result of current conventional materials not invoking suitable cellular responses to promote adequate osteointegration to enable implanted devices to be successful for long periods (Balasundaram and Webster, 2006; Barrère et al., 2008). Metallic materials are normally used for load-bearing members such as pins and plates, femoral stems, etc. Ceramics, such as alumina and zirconia, are used for wear applications in joint replacements, while hydroxyapatite is used for bone bonding applications to assist implant integration. Polymers, such as ultra high molecular weight polyethylene (UHMWP), are used as articulating surfaces against ceramic components in joint replacements. Porous alumina has also been used as a bone spacer to replace large sections of bone which have had to be removed due to disease (www.azom, 2004; http://academic.uprm.edu). In applications involving the loading phase, the best material has been titanium and its alloys, whereas calcium phosphate seems to be the best material to be used in joint replacement and osseointegration (the degree to which bone will grow next to or integrate into the implant). Titanium is used primarily for the loading faces, which include the pin structure, and the fabrication of plates and femoral stems. The integration of a biomaterial to bone involves, essentially, two processes: interlocking with bone tissue and chemical interactions with bone constituents. The direct bonding of orthopedic biomaterials with collagen is rarely considered; however, several non-collagenic proteins have been shown to adhere to biomaterial surfaces (Rey, 1998). Many studies have been reported on the biointegration of orthopedic devices. Hydroxyapatite (HA) films have been widely recognized for their biocompatibility and utility in promoting biointegration of implants in both osseous and soft tissue. In a study on hydroxyapatite-coated (by electroplating) cp-titanium implants, Badr and El Hadary (2007) showed the formation of recognizable osseointegration of bone regeneration with more and denser bone trabeculae, and concluded that electroplating provided a thin

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Biointegration: an introduction

3

and uniform pure crystalline hydroxyapatite coating. The characterization of the precipitated film is promising for clinically successful long-term bone fixation. An AFM analysis of roughness on seven materials widely used in bone reconstruction carried out by Covani et al. showed that the biointegration properties of bioactive glasses can also give an answer in terms of surface structures in which chemical composition can influence directly the biological system (e.g. with chemical exchanges and development of specific surface electrical charge) and indirectly, via the properties induced on tribological behavior that expresses itself during the smoothing of the surfaces (Covani et al., 2007). The biological behavior of an implant, such as osseointegration, depends on both the chemical composition and the morphology of the surface of the implant. Irradiation with laser light (Nd:YAG (λ = 1064 nm, τ = 100 ns)) is used for the surface modification of Ti-6Al-4V – which is widely used in implantation to enhance biointegration (Mirhosseini et al., 2007). Conventional sputtering techniques have shown some advantages over the commercially available plasma spraying method for generating hydroxyapatite (HA) films on metallic substrates; however, the as-sputtered films are usually amorphous, which can cause some serious adhesion problems when post-deposition heat treatment is necessitated. Nearly stoichiometric, highly crystalline HA films strongly bound to the substrate were obtained by an opposing radio frequency (RF) magnetron sputtering approach. HA films have been widely recognized for their biocompatibility and utility in promoting biointegration of implants in both osseous and soft tissue (Hong et al., 2007). Oudadesse et al. (2007) studied the in vitro behavior of compounds in contact with simulated body fluid (SBF) and in vivo experiments in a rabbit’s thigh bones. The inductively coupled plasma–optical emission spectroscopy (ICP-OES) method was used to study the eventual release of Al from composites to SBF and to evaluate the chemical stability of composites characterized by the succession of SiO4 and AlO4 tetrahedra. The results obtained show the chemical stability of composites. At the bone– implant interface, the intimate links revealed the high quality of the biointegration and the bioconsolidation between composites and bony matrix. Histological studies confirmed good bony bonding and highlighted the total absence of inflammation or fibrous tissues, indicating good biointegration (Oudadesse et al., 2007). Zanotti and Verlicchi (2006) proposed a bioglass– alumina spacer that could perform an excellent arthrodesis by a mechanical stabilization (primary one) and a biological bio-mimetic stabilization (biointeraction, biointegration and biostimulation). Other advantages are easy use, even in the low somatic interspaces (C7-D1 type), reduction of convalescence, and reduced costs of this type of device. That makes it interesting, even in societies with low economic well-being (Zanotti and Verlicchi, 2006).

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Biointegration of medical implant materials

The process by which these materials are integrated into the organspecific repair cascade is named ‘bone remodeling’, which is the concerted interplay of two cellular activities: osteoclastic bone resorption and osteoblastic bone formation. The latter physiologic process not only maintains bone mass, skeletal integrity and skeletal function but is also the cellular process that determines structural and functional integration of bone substitutes. A molecular understanding of this process is therefore of paramount importance for almost all aspects of the body’s reaction to biomaterials and might help us to understand, at least in part, the success or the failure of various materials used as bone substitutes. Recent genetic studies have demonstrated that there is no tight cross-control of bone formation and bone resorption in vivo and that there is also a central axis controlling bone remodeling, radically enhancing our understanding of this process. Amling et al. (2006) have shown how an understanding of bone remodeling – the physiologic process required for the biologization of bone substitutes – has evolved during recent years, providing a platform for the design, development and application of improved biomaterials. Roughness was evaluated by measuring root mean square (RMS) values and RMS/average height (AH) ratio, in different dimensional ranges, varying from 100 microns square to a few hundreds of nanometers. The results showed that titanium presented a lower roughness than the other materials analyzed, frequently reaching statistical significance (Covani et al., 2007). Conversely, bioactive materials such as hydroxyapatite (HA) and bioactive glasses have demonstrated an overall higher roughness. In particular, this study focuses attention on AP40 and especially RKKP, which proved to have a significant higher roughness at low dimensional ranges. This determines a large increase in surface area, which is strongly connected with osteoblast adhesion and growth, and also with protein absorption (Fig. 1.1). One should mention here the famous osseointegration concept evolved by Per Ingvar Brånemark, closely coupled with the design of a cylindrical titanium screw (Fig. 1.2) (Albrektsson et al., 1981; Brånemark et al., 1985) having a specific surface treatment to enhance its bioacceptance (Adell et al., 1990). The titanium screw (Fig. 1.2) underwent many animal and, subsequently, human clinical trials to test the success rate, the concept and the design of this implant. A fixture is osseointegrated if it provides a stable and apparently immobile support of a prosthesis under functional loads, without pain, inflammation, or loosening. Titanium’s ability to be integrated in the bone has been known for more than 25 years of experience and research that form the basis of the knowledge and use of implant technology today. Osseointegration of an implant is a direct structural and functionally stable connection

© Woodhead Publishing Limited, 2010

Biointegration: an introduction Titanium

Bone

Osseointegrated (a)

Titanium Connective tissue

5 Bone

Non-integrated (b)

1.1 (a) and (b) Biointegration of titanium showing a lower roughness (reproduced from Covani et al., 2007 with permission from Global Rights, John Wiley & Sons Inc.).

1.2 The titanium screw (reproduced with permission from Brånemark et al., Tissue-integrated Prostheses: Osseointegration in Clinical Dentistry, Quintessence Publishing Co., Chicago, 1985).

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Biointegration of medical implant materials

between the living bone and the surface of an implant that is exposed to mechanical load (http://www.dental-oracle.org). Specialized coatings are being developed for orthopedic implants, biomedical use, and other scientific applications. A HA coating for surface improvement gave rapid osseointegration and biointegration in four weeks with 90% of implant–bone contact at ten months in contrast to titanium alone, which required ten weeks to be osseointegrated with 50% implant– bone contact at ten months. A very recent material is osteopontin, an extracellular glycosylated bone phosphoprotein with a polypeptide backbone that makes dead metal ‘come alive’. Surrounding cells ‘don’t see an inert piece of metal, they see a protein and it’s a protein they know’. Other novel innovations include the development of nanostructure materials and diagnostic techniques for both in vitro and in vivo applications (Namavar et al., 2007). Typically, in regards to orthopedic devices, the primary concerns are wear, infection and failure of biointegration (Harris and Richards, 2006; Viceconti et al., 2004). The survival rate of an implant under optimal conditions is at least 96% after five years. Zim reports development of a high-porosity expanded polytetrafluoroethylene that has been fabricated to provide a softer feel with less shrinkage and migration because of better biointegration and cellular ingrowth (Zim, 2004). Long-term results with porous polyethylene have demonstrated superior biocompatibility and minimal complications. Hydroxyapatite cement has been associated with an immunoguided delayed inflammatory reaction that leads to thinning of the overlying skin and exposure of the implant. Applications of distraction osteogenesis are rapidly expanding and include deformities of the mandible, midface, and cranium. There has been a trend toward the use of internal hardware, and internal devices are being developed to deliver a greater degree of vector control. Biodegradable devices have been developed to eliminate the second surgical procedure necessary for hardware removal. In the future, successful tissue engineering could eliminate many of the drawbacks associated with implants and osteotomies. The ability to stimulate stem cells to generate autogenous bone has been demonstrated. Computer technology has been successfully used to integrate laser surface scanning and digitizing with computer-aided design and manufacturing to produce facial prostheses. Technologic advances in biomaterials, distraction hardware, computer modeling, and tissue engineering will continue to supply the surgeon’s repertoire with improved methods to augment and restore the craniomaxillofacial skeleton. Whenever metallic devices are implanted in vivo, successful biointegration requires that host cells colonize the highly reactive implant surface (Schmidt and Swiontkowski, 2008). Bacteria such as staphylococci can also become adherent to metallic or polymeric implants and will compete with

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Biointegration: an introduction

7

host cells for colonization of the implant surface. It has been demonstrated in animal models that contaminated fractures without internal fixation develop clinical infection more commonly than similar fractures treated with internal fixation at the time of colonization. Because of the potential for infection whenever internal fixation is utilized, appropriate prophylactic antibiotic coverage for staphylococci and Gram-negative organisms should be provided.

1.3

Biointegration of biomaterials for dental applications

The biomaterials community is producing new and improved implant materials and techniques to meet the growing demands for biomaterials in dental applications. The main property required of a biomaterial is that it does not illicit an adverse reaction when placed into service (www.azom.com). Thus, a variety of materials – metallic (pins for anchoring tooth implants and as parts of orthodontic devices), ceramics (tooth implants including alumina and dental porcelains) and polymeric (orthopedic devices such as plates and dentures) – with excellent biointegration have been developed and have been placed into service. Starting with the ‘Integral Biointegrated Dental Implant System’, consisting of a titanium implant cylinder coated with calcitite that permitted the bone to actually bond with the implant surface in a jaw restoration (Roling, 1989), dental surgery has undergone a revolution in both implant techniques and materials technology. It was shown early on that the surface oxide of titanium appears to be central to the ability of titanium implants to achieve osseointegration, and ceramic coatings appear to improve the ingrowth of bone and promote chemical integration of the implant with the bone (Wataha, 1996). Badr and El Hadary (2007) have reported development of osseointegration of the regenerated bone when hydroxyapatite was coated onto the surface of commercially pure titanium (cpTi) implants using an electroplating technique. Pelsoczi et al. obtained more effective osseointegration on surface modifications of titanium implants with an excimer laser. In this case, it was easy to achieve the desired morphology (microstructure) and physical-chemical properties that control the biointegration process (Pelsoczi et al., 2004). X-ray photoelectron spectroscopy (XPS) studies show that laser treatment, in addition to micro-structural and morphological modification, results in a decrease of surface contamination and thickening of the oxide layer. Thin film deposition of ceramic oxides onto titanium by excimer lasers and pulsed lasers has been successfully employed by other groups to improve the surface characteristics for facilitating biointegration, e.g. pulsed laser deposition of bioceramic thin films from human teeth (Smausz et al., 2004) and surface modifications induced

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Biointegration of medical implant materials

by ns and sub-ps excimer laser pulses on titanium implant material (Bereznai et al., 2003).

1.4

AlphaCor artificial corneal experience

AlphaCorTM is a biocompatible, flexible, one-piece artificial cornea (keratoprosthesis) designed to replace a scarred or diseased native cornea. It is a one piece convex disc consisting of a central transparent optic and an outer skirt that is entirely manufactured from poly(2-hydroxyethyl methacrylate) or PHEMA. AlphaCor’s material and patented features are designed to promote retention and optimize patient outcome. The outer skirt is an opaque, high water-content, PHEMA sponge. The porosity of the sponge encourages biointegration with host tissue and thus promotes retention of the implanted device. The central optic core is a transparent PHEMA gel, providing a refractive power similar to that of the human cornea. The optic core is designed to allow the patients’ visual potential to be achieved. The junctional zone between the skirt and central optic is the interpenetrating polymer network or IPN. This is a permanent bond formed at the molecular level and is designed to prevent the down-growth of cells around the optic, which can lead to the formation of retroprosthetic membranes, one of the major complications historically associated with artificial corneas (http://www.medcompare.com; http://www.pricevisiongroup.com). The World Health Organization (WHO) reports that corneal blindness affects more than 10 million people worldwide; however, only 100 000 people received corneal transplants each year. This shortfall is due to a combination of inadequate supply of donor corneas and the unsuitability of some patients to receive a corneal graft. AlphaCor is designed for use in patients who have had multiple failed corneal transplants or in those patients in whom a donor graft is likely to fail. Its patented design features are aimed to promote retention, minimize post-operative complications and restore vision in patients who cannot receive, or are unlikely to have, a beneficial outcome from a human donor graft. AlphaCor is available in two versions, to suit those with a natural lens (phakic) or artificial lens (pseudophakic) and for those without a lens (aphakic). Keratoprosthesis for artificial cornea surgery is a procedure for restoring the sight of patients suffering from a severely damaged anterior segment due to trauma, chemical burns, infections, etc. The ideal keratoprosthesis would be inert and not be rejected by the patient’s immune system, be inexpensive, and maintain long-term clarity. In addition, it would be quick to implant, easy to examine, and allow an excellent view of the retina. Coassin et al. (2007) reported histopathologic and immunologic characteristics of late artificial corneal failure in a small series of patients who

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Biointegration: an introduction

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1.3 AlphaCorTM artificial cornea (reproduced with permission from M/s Addition Technology Inc, 950 Lee Street, Des Plaines, IL 60016, USA).

underwent AlphaCor implantation, but light microscopic examination of the specimens disclosed adequate biointegration with no foreign body response. Immunofluorescence studies of the skirt exhibited expression of inflammatory cytokines such as interleukin-1β (IL-1β) and tumor necrosis factor α (TNF-α), and some interferon γ (IFN-γ). The keratocytes stained positively for Thy-1 and smooth muscle actin, but negatively for CD34. The AlphaCor implant (Fig. 1.3) is a viable method of treatment for multiple failed PKPs, but it may be associated with unique complications, including corneal stromal melting, focal calcification, and retroprosthetic membrane formation. Infectious keratitis may be a risk factor for corneal stromal melting and needs to be managed aggressively. Explantation of the implant is essential if the skirt is exposed (Chow et al., 2007). Hicks et al. (2005) showed that histologic findings of the AlphaCor skirt in humans are consistent with earlier animal studies. Their study confirmed that biointegration by host fibroblastic cells, with collagen deposition, occurred after AlphaCor implantation in humans. In cases in which stromal melting had occurred, biointegration was seen to be reduced. On correlating preoperative clinical factors with biointegration observed histologically, preoperative vascularization appears not to be required for AlphaCor biointegration (Hicks et al., 2005). Another study demonstrated that systemic factors

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affected the risk of retroprosthetic membrane formation with AlphaCor (Hicks and Hamilton, 2005). Hicks et al. (2005) further showed that device biointegration gets reduced in the cases of patients with a history of ocular herpes simplex virus (HSV) because the extensive lamellar corneal surgery involved in AlphaCor implantation may precipitate reactivation of latent HSV, such that reactivation and resultant inflammation could facilitate melting of corneal stromal tissue anterior to the device (Hicks et al. 2002). Chirila (2006), in a recent publication, claims that the first keratoprosthesis based on polyurethane was made in 1985 by Lawrence Hirst, an Australian ophthalmologist then working in St Louis, USA. This keratoprosthesis, which also had a porous skirt, was inserted intralamellarly in a monkey cornea and was followed up clinically for about three months. There were no significant postoperative complications, and the histology of the explant indicated proper biointegration of the prosthetic skirt within the host stromal tissue (Chirila, 2006; Chirila et al., 1998). Hydrogel lenses may even make their way deeper into the eye, as replacements for inner-eye lenses damaged by cataracts.

1.5

Biointegration and functionality of tissue engineering devices

Experiments on animals have underlined the importance of vascularization for biointegration and functionality of any given tissue engineering device. Polykandriotis et al. recently showed that the presence of a vascular bed prior to cell transplantation might protect against hypoxia-induced cellular death, especially at central portions of the matrix, and therefore ensure physiological function of the device. The generation of vascularized bioartificial tissue substitutes might offer new modalities of surgical reconstruction for use in reparative medicine (Polykandriotis et al., 2006).

1.6

Percutaneous devices

Percutaneous devices play an essential role in medicine; however, they are often associated with a significant risk of infection. One approach to circumvent infection would be to heal the wound around the devices by promoting skin cell attachment (Fukano et al., 2006). Biointegration through human fibronectin (FN) plays a key role in the biointegration of implants, as the success depends on adsorption of proteins like FN. Indeed FN can be an intermediary between the biomaterial surface and cells (Sousa et al., 2005). Isenhath et al. (2007) developed an in vivo model that permits examination of the implant/skin interface and that will be useful for future studies designed to facilitate skin cell

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attachment where percutaneous devices penetrate the skin. The space created between the skin and the device becomes a haven for bacterial invasion and biofilm formation, and this results in infection. Sealing this space via integration of the skin into the device is expected to create a barrier against bacterial invasion. Porous poly(2-hydroxyethyl methacrylate) (PHEMA) rods were implanted for seven days in the dorsal skin of C57 BL/6 mice. The porous PHEMA rods were surface-modified with carbonyldiimidazole (CDI) or CDI plus laminin 5, with unmodified rods serving as control. Implant sites were sealed with 2-octyl cyanoacrylate; corn pads and adhesive dressings were tested for stabilization of implants. All rods remained intact for the duration of the study. There was histological evidence of both epidermal and dermal integration into all PHEMA rods, regardless of treatment. The effects of hyperbaric oxygen (HBO) therapy on biointegration of porous polyethylene (PE) implanted beneath dorsal burn scar and normal skin of Sprague-Dawley rats were microscopically examined and the ratio of fibrovascular ingrowth (FVI) was determined for each rat (Dinar et al., 2008). The results showed that HBO therapy enhanced biointegration of porous PE in hypoxic burn scar areas via improving collagen synthesis and neovascularization; otherwise, it apparently delayed tissue ingrowth into a porous structure implanted in normal healthy tissues.

1.7

Future trends

The future of biointegration and the future of implants are considered to be bright, as advancements in frontier biomaterials are advancing rapidly to unravel the physiologic process required for the biologization of the implants and developing materials that become intrinsically integrated into the organ-specific repair mechanisms. One example is the ‘designer implant’, which could carry different types of proteins, one set to spur soft tissue healing, another to encourage hard tissue growth on another front. According to Rush, future devolvement of osteobiologic materials will no doubt replace materials currently being used (Rush, 2005). Techniques to improve biointegration and manipulation of the healing environment will be developed such that future graft substitutes may exceed even autogenous bone in their reliability. An understanding of the cascade of events that occurs with bone healing and graft incorporation will enhance the chances to augment or manipulate the grafting process. The biomaterials community is producing new and improved implant materials and techniques to meet this demand. A counter force to this technological push is the increasing level of regulation and the threat of litigation. To meet these conflicting needs it is necessary to have reliable methods of characterization of the material and material/host tissue interactions (www.azom.com). In addition,

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progress on the road to regenerating major body parts, salamander-style, could transform the treatment of amputations and major wounds (Muneoka et al., 2008). At the same time, the newly emerging nanobiotechnology is revolutionizing capability to resolve biological and medical problems by developing subtle biomimetic techniques. The use of nanoscale materials is expected to increase dramatically in many applications of medicine and surgery. It is interesting to note that the size of these nanomaterials is comparable to many biological systems and so there is large scope for biointegration. Additionally, nanomaterials exhibit fundamentally different properties from their properties in bulk, such that they can be tailor-made to provide properties to suit a specific application (Namavar, 2003; Christenson et al., 2007). Significant advancements are expected to achieve the desired goals and their clinical use, especially in areas such as nanostructured coatings, nanostructured porous scaffoldings and other nanobiomaterials. The current trends in nanobiotechnology, thus, offer a bright future through the use of nanobiomaterial in achieving biointegration.

1.8

References

adell r, eriksson b, lekholm u, brånemark p i and jemt t (1990), ‘Long-term follow-up study of osseointegrated implants in the treatment of totally edentulous jaws’, Int J Oral Max Impl, 5, 347–359. albrektsson t, brånemark p i, hansson h a, lindstrom j (1981), Osseointegrated titanium implants: Requirements for ensuring a long-lasting, direct bone-toimplant anchorage in man. Acta Orthop Scand, 52, 155–170. amling m, schilling a f, pogoda p, priemel m and rueger j m (2006), ‘Biomaterials and bone remodeling: The physiologic process required for biologization of bone substitutes’, Eur J Trauma, 32, 102–106. Doi: 10.1007/s00068-006-6049-6. badr n a and el hadary a a (2007), ‘Hydroxyapatite-electroplated cp-titanium implant and its bone integration potentiality: An in vivo study’, Impl Dent, 16, 297–308. Doi: 10.1097/ID.0b013e31805d7dc4. balasundaram g and webster t j (2006), ‘Nanotechnology and biomaterials for orthopedic medical applications’, Nanomedicine, London, 1, 169–176. barrère f, mahmood t a, de groot k and van blitterswijk c a (2008), ‘Advanced biomaterials for skeletal tissue regeneration: Instructive and smart functions’, Mat Sci Eng R, 59, 38–71. bereznai m, pelsöczi i, tóth z, turzó k, radnai m, bor z and fazekas a (2003), ‘Surface modifications induced by ns and sub-ps excimer laser pulses on titanium implant material’, Biomaterials, 24, 4197–4203. Doi: 10.1016/ S0142-9612(03)00318-1. brånemark p i, zarb g a and albrektsson t (1985) Tissue-integrated Prostheses: Osseointegration in Clinical Dentistry, Chicago, Quintessence Publishing Co., pp. 1–356.

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chirila t v, hicks c r, dalton p d, vijayasekaran s, lou x, hong y, clayton a b, ziegelaar b w, fitton j h, platten s, crawford g j and constable i j (1998), ‘Artificial cornea’, Prog Polym Sci, 23, 447–473. chirila t v (2006), ‘First development of a polyurethane keratoprosthesis and its Australian connection: An unbeknown episode in the history of artificial cornea’, Clin Exp Ophthalmol, 34, 485–488. Doi: 10.1111/j.1442-9071.2006. 01251.x. chow c c, kulkarni a d, albert d m, darlington j k and hardten d r (2007), ‘Clinicopathologic correlation of explanted AlphaCor artificial cornea after exposure of implant’, Cornea, 26, 1004–1007. Doi: 10.1097/ICO.0b013e3180e799f0. christenson e m, anseth k s, van den beucken j j j p, chan c k, ercan b, jansen j a, laurencin c t and mikos a g (2007), ‘Nanobiomaterial applications in orthopedics’, J Orthopaed Res, 25, 11–22. coassin m, zhang c, green w r, aquavella j v and akpek e k (2007), ‘Histopathologic and Immunologic Aspects of AlphaCor Artificial Corneal Failure’, Am J Ophthalmol, 144, 699–704 e4. Doi: 10.1016/j.ajo.2007.07.025. cochran d (1996), ‘Implant therapy I’, Ann Periodontol, 1, 707–791. covani u, giacomelli l, krajewski a, ravaglioli a, spotorno l, loria p, das s and nicolini c (2007), ‘Biomaterials for orthopedics: A roughness analysis by atomic force microscopy’, J Biomed Mater Res A, 82, 723–730. dinar s, agir h, sen c, yazir y, dalcik h and unal c (2008), ‘Effects of hyperbaric oxygen therapy on fibrovascular ingrowth in porous polyethylene blocks implanted under burn scar tissue: An experimental study’, Burns, 34, 467–473. Doi: 10.1016/j.burns.2007.04.014. fukano y, knowles n g, usui m l, underwood r a, hauch k d, marshall a j, ratner b d, giacelli c, carter w g, fleckman p and olerud j e (2006), ‘Characterization of an in vitro model for evaluating the interface between skin and percutaneous biomaterials’, Wound Repair Regen, 14, 484–491. Doi: 10.1111/j.1743-6109.2006. 00138.x. harris l g and richards r (2006), ‘Staphylococci and implant surfaces: A review’, Injury, 37, S3–S14. Doi:10.1016/j.injury.2006.04.003. hicks c r, crawford g j, tan d t, snibson g r, gondhowiardjo t d, lam d s c and downie n (2002), ‘Outcomes of implantation of an artificial cornea, AlphaCor: Effects of prior ocular herpes simplex infection’, Cornea, 21, 685–690. Doi: 10.1097/00003226-200210000-00010. hicks c r and hamilton s (2005), ‘Retroprosthetic membranes in AlphaCor patients: Risk factors and prevention’, Cornea, 24, 692–698. Doi: 10.1097/01.ico. 0000154380.13237.ea. hicks c r, werner l, vijayasekaran s, mamalis n and apple d j (2005), ‘Histology of AlphaCor skirts: Evaluation of biointegration’, Cornea, 24, 933–940. Doi: 10.1097/01.ico.0000160969.50706.7f. hong z, luan l, paik s b, deng b, ellis d e, ketterson j b, mello a, eon j g, terra j and rossi a (2007), ‘Crystalline hydroxyapatite thin films produced at room temperature – An opposing radio frequency magnetron sputtering approach’, Thin Solid Films, 515, 6773–6780. Doi: 10.1016/j.tsf.2007.02.089. http://academic.uprm.edu/~mgoyal/materialsmay2004/k04orthopaedic.pdf. http://www.dental-oracle.org/uk/implant/Pages/c.html. http://www.karger.com/gazette/65/lidgren/art_5_0.html.

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http://www.medcompare.com/details/40895/AlphaCorAndtrade-Artificial-Cornea. html. http://www.pricevisiongroup.com/cornea_transplant.html. isenhath s n, fukano y, usui m l, underwood r a, irvin c a, marshall a j, hauch k d, ratner b d, fleckman p and olerud j e (2007), ‘A mouse model to evaluate the interface between skin and a percutaneous device’, J Biomed Mater Res Part A, 83, 915–922. Doi: 10.1002/jbm.a.31391. meffert r m, block m s and kent j n (1987), ‘What is osseointegration?’, Int J Periodontics Restorative Dent, 7, 9–21. mirhosseini n, crouse p l, schmidth m j j, li l and garrod d (2007), ‘Laser surface micro-texturing of Ti-6Al-4V substrates for improved cell integration’, Appl Surf Sci, 253, 7738–7743. muneoka k, han m and gardiner d m (2008), ‘Regrowing human limbs’, Sci Amer 298, 56–63. namavar f, jackson j d, sharp g, mann e e, bayles k, cheung c l b, feschuk c a, varma s, haider h and garvin k l (2007), Searching for Smart Durable Coatings to Promote Bone Marrow Stromal Cell Growth While Preventing Biofilm Formation, Mater Res. Soc. Symp Proc., Vol. 954 © Materials Research Society 0954-H04-04. namavar f (2003), ‘Applications of nanotechnology for alternative bearing surfaces in orthopaedics’, Proceedings of the 8th Ceramics, Cells and Tissues MeetingSeminar, Faenza, Italy March, 2003, Volume edited by A Ravaglioli and A Krajewski, ISTEC-CNR (December 2003). oudadesse h, derrien a c, mami m, martin s, cathelineau g and yahia l (2007), ‘Alumino silicates and biphasic HA-TCP composites: Studies of properties for bony filling’, Biomed Mater, 2, art.no.S09, S59–S64. Doi: 10.1088/1748-6041/2/1/ S09. pelsoczi k i, bereznai m, tóth z, turzó k, radnai m, bor z and fazekas a (2004), ‘Surface modifications of titanium implant material with excimer laser for more effective osseointegration’ (Titán-minták felületének módosítása excimer lézerrel a hatékonyabb osszeointegráció erdekében), Fogorvosi Szemle, 97, 231–237. polykandriotis e, arkudas a, euler s, beier j p, horch r e and kneser u (2006), ‘Prevascularisation strategies in tissue engineering’ (Prävaskularisationsstrategien im tissue engineering), Handchirurgie Mikrochirurgie Plastische Chirurgie, 38, 217–223. rey c (1988), ‘Orthopedic biomaterials, bioactivity, biodegradation: A physicalchemical approach’, J Biomech, 31, Supplement 1, July 1998, 182. roling t (1989), ‘Biointegration revolutionizes dental surgery’, Sulzer Technical Review, 71, 7–10. rush s m (2005), ‘Bone graft substitutes: Osteobiologics’, Clinics in Podiatric Medicine and Surgery, 22, 619–630. Doi: 10.1016/j.cpm.2005.07.004. schmidt a h and swiontkowski m f (2008), ‘Pathophysiology of infections after internal fixation of fractures’, J Am Acad Orthop Surg, 285–291. smausz t, hopp b, huszár h, töth z and kecskeméti g (2004), ‘Pulsed laser deposition of bioceramic thin films from human tooth’, Appl Phys A-Mater, 79, 1101–1103. sousa s r, moradas-ferreira p and barbosa m a (2005), ‘TiO2 type influences fibronectin adsorption’, J Mater Sci-Mater M, 16, 1173–1178. Doi: 10.1007/ s10856-005-4725-4.

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viceconti m, davinelli m, taddei f and cappello a (2004), Automatic generation of accurate subject-specific bone finite element models to be used in clinical studies, J Biomech, 37, 1597–1605. Doi:10.1016/j.jbiomech.2003.12.030. wataha j c (1996), ‘Materials for endosseous dental implants’, J Oral Rehabil, 23, 79–90. williams d f (1999), ‘The Williams Dictionary of Biomaterials’, Liverpool University Press, Liverpool. www.azom.com, Biomaterials: an overview. zanotti b and verlicchi a (2006), ‘Is one cervical prosthesis equal to another?’ (Una protesi cervicale vale l’altra?), Rivista Medica, 12, 125–130. zim s (2004), ‘Skeletal volume enhancement: Implants and osteotomies’, Current Opinion in Otolaryngology and Head and Neck Surgery, 12, 349–356. Doi: 10.1097/01.moo.0000130576.04818.55.

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2 Biocompatibility of engineered soft tissue created by stem cells P. A. C L A R K, University of Wisconsin–Madison, USA; and J. J. M AO, Columbia University, USA

Abstract: Orthopedic and dental implants replace millions of arthritic, traumatic or lost skeletal or dental structures. All orthopedic or dental implants can fail, and the central reason for failure is that metallic implants do not remodel with host tissue. Current implants rely primarily on tissue growth onto implants or an ‘outside-in’ strategy. This chapter discusses an ‘inside-out’ strategy to induce tissue ingrowth by cytokine delivery. Drug-eluting porous implants have the advantage not only of reducing bulk metal mass, but also of harboring cytokines that are programmed to release into surrounding tissue. This coupled inside-out and outside-in strategy improves bone ingrowth. Key words: orthopedic implants, medical implants, cardiac implants, dental implants, growth factors, cytokines, controlled release, microencapsulation.

2.1

Introduction

Tissue and organ defects resulting from trauma, chronic diseases, tumor resection or congenital anomalies necessitate the restoration of the lost anatomical structures. Due to a lack of biological replacements for skeletal structures, implantation of biocompatible metals such as titanium is currently the preferred treatment (Ratner et al., 1996; Misch, 1993; Kienapfel et al., 1999). Despite high success rates of initial anchorage (over 90%) (Ashley et al., 2003), titanium implants require long healing times before functional loading, and are subject to failure from inadequate initial bone ingrowth or long-term osteolysis at the bone–implant interface. Current approaches in modifying titanium implants to overcome these limitations focus on biomaterial composition and processing, surface roughening, and chemical surface modification, among others. Taking cues from biology and tissue engineering has led to the idea of biointegration, entailing the use of biologically active agents to modulate the bone ingrowth process and improve implant anchorage. Biointegration of orthopedic implants represents a daunting task considering the complex environment of healing and homeostatic bone, but fortunately the fields of tissue engineering and drug 19 © Woodhead Publishing Limited, 2010

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delivery have developed micron- and nano-scale systems for controlled release of various biologically active agents. This chapter will first address the key biological considerations of biointegration of orthopedic implants. Mimicking these processes using drug delivery systems toward improved short- and long-term efficacy in these implants will then be discussed, concluding with future trends in this emerging field.

2.2

Bone: from tissue to molecular organization

Bone is the main weight-bearing tissue of the body, with varying and complex macroscopic designs due to its distinct functions in different regions of the body. In general, bone exists either as cortical with low porosity and high density, or cancellous (or trabecular) with microscopic interconnecting bony trabecula giving macroscopically high porosity and low density (Marks and Odgren, 2002). Biochemically, bone is composed of about 35% organic matrix (osteoid), mainly Type I collagen fibers along with proteoglycans and noncollagenous proteins, and about 65% inorganic mineral, mainly calcium and phosphate in the form of hydroxyapatite (Lind, 1998; Misch, 1993). This general composition gives bone marked rigidity while retaining some elasticity (Marks and Odgren, 2002), with the collagen fibers of the organic matrix providing high tensile strength to resist pulling forces and the inorganic mineral providing high compressive strength to resist crushing forces (Marks and Odgren, 2002; Misch, 1993; Alberts et al., 2002). A key facet of bone tissue during development and maintenance is the constant re-organization of the extracellular matrix to satisfy local loadbearing requirements. This process is driven by the two main cell phenotypes of bone: bone-forming osteoblasts and bone-resorbing osteoclasts (Marks and Odgren, 2002; Hole and Koos, 1994). Acting as possible sensors and signaling agents for the osteoblasts and osteoclasts are the osteocytes, post-mitotic terminally differentiated osteoblasts encased in bone matrix that communicate via long processes known as canaliculi (Marks and Odgren, 2002). Located in the cavities of long bones and among trabecula in cancellous bone is the bone marrow. This tissue contains both red marrow, the site of new blood cell production or hematopoiesis throughout life, and yellow marrow, which is mostly fat cells (Hole and Koos, 1994). The bone marrow generally transitions from red to yellow with age, although this trend can be reversed in injurious or other special instances (Hole and Koos, 1994). The marrow contains a milieu of cells, including red and white blood cells, osteoblasts, fibroblasts, adipocytes (fat cells), and blood vessel cells (Hole and Koos, 1994). Fibroblast-like cells residing within the bone marrow stroma, or connective tissue of the marrow, have also been isolated that

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possess extensive proliferative potential and differentiation ability to multiple mesenchymal lineages, including osteoblasts, chondrocytes (cartilage cells), and adipocytes (Pittenger et al., 1999; Caplan, 1991a,b; Alhadlaq and Mao, 2004; Cassiede et al., 1996). These multipotential cells, termed mesenchymal stem cells (MSCs) or bone marrow stromal cells, likely play a key role in repair after injury, and are present throughout life (Caplan, 1991b; Alhadlaq and Mao, 2004). Lining the outer wall of the marrow cavity and the outer surface of bones are thin linings of tissue called the endosteum and periosteum, respectively. These tissues are similar in composition and morphology, being composed of flattened cells (Marks and Odgren, 2002). Recently, the endosteal lining has been identified as an important hematopoietic stem cell (HSC) ‘niche’, the specialized compartment in which stem cells reside (Scadden, 2006; Taichman, 2005). Through strict control of the microenvironment, the endosteal cells maintain the HSCs, which can differentiate to every blood lineage, until they are needed (Taichman, 2005; Scadden, 2006). The endosteum and periosteum may also contain osteoprogenitor cells that can mobilize after injury (Hutmacher and Sittinger, 2003; Hanada et al., 2001). Not to be forgotten, like any tissue, bone and its marrow require a rich vascular supply for oxygen and nutrients and for disposal of waste products. As will be discussed later in the chapter, vessel formation or ingrowth is critical for bone formation during development and after injury. Mural cells associated with blood vessels, particularly the pericytes, have demonstrated multilineage potential and may also participate in bone repair after injury (Collett and Canfield, 2005; Doherty and Canfield, 1999). The tissue and cellular processes that organize and maintain bone are molecularly coordinated and controlled largely by bioactive chemicals termed cytokines or growth factors (Gilbert, 1997). During development, homeostasis, and after injury, a multitude of skeletal growth factors act as both temporal and spatial coordinating molecules to induce chemotaxis, mitosis, differentiation, changes in extracellular matrix production, and even apoptosis (Roberts, 2000; Alliston and Derynck, 2000). These cytokines can exert their effects on local cells (paracrine), on the same cells that released them (autocrine), or after absorption and transport via the bloodstream (endocrine) (Lind, 1998). Some cytokines exert their effects on very specific lineages of cells while others exhibit context-dependent effects on multiple cell phenotypes. Growth factor effects on cells depend heavily on the dosage, with most cells demonstrating biphasic responses. Some of the most well-studied cytokines in skeletal biology are members of the transforming growth factor β (TGFβ) superfamily, which include multiple TGFβ isoforms and the bone morphogenetic proteins (BMPs). Members of this superfamily play important and oftentimes critical roles in the growth and

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maintenance of bone and cartilage. Perhaps most strikingly, the BMPs were discovered by Marshall Urist in 1965 by their ability to induce ectopic bone formation in skeletal muscle (Lou, 2001; Linkhart et al., 1996; Lind, 1998; Urist, 1965; Urist et al., 1979). Fibroblast growth factors (FGFs) and the insulin-like growth factors (IGFs) also participate in many skeletal development and repair processes (Marie et al., 2002; Linkhart et al., 1996). Important to vasculogenic and angiogenic processes, and therefore to bone development and homeostasis, are vascular endothelial growth factor (VEGF) and platelet-derived growth factor (PDGF) (Lind, 1998; Franceschi, 2005; Gerstenfeld et al., 2003). Details of these, as well as other cytokines, in skeletal biology will be expounded upon later in the chapter, but their early introduction is warranted as their appearance is a common theme throughout skeletal development, homeostasis, and repair.

2.3

Bone development

Becoming apparent in recent years is the uncanny resemblance that repair processes have to embryonic development (Gerstenfeld et al., 2003; Caplan, 2003), and therefore it is important to confer at least a general understanding of the mechanisms of bone formation during embryogenesis. Embryonic limb bud formation begins by budding from the lateral surface at specific levels (Tuan, 2004). Undifferentiated mesenchyme begins to pocket out as a limb field, expressing FGF10. Thickening of the edge of the bud forms the apical ectoderm ridge, and begins to express FGF8, BMP2, BMP4, and Msx2. At this point, spatial patterning (ventral vs. dorsal, anterior vs. posterior) is initiated, as determined by gradients of Hox genes as well as by Wnt (related to the developmental molecules wingless in Drosophila) and FGF4. The mesenchyme then condenses, with these cells secreting a variety of signaling factors, including growth and differentiation factor (GDF) 5, BMP2, BMP4, BMP7, and FGF9 (Tuan, 2004; Ornitz, 2005). Condensation of mesenchymal cells to trigger lineage specification is common throughout development of bone and cartilage, suggesting critical roles for extracellular matrix (ECM) components (Gilbert, 1997) and cell shape. Lineage specification of adult MSCs isolated and expanded ex vivo also require proper cell shape and ECM (Cassiede et al., 1996; Pittenger et al., 1999; McBeath et al., 2004), exemplifying the importance of basic developmental concepts in regenerative medicine. Condensing mesenchymal cells begin to differentiate to chondrocytes and express noggin, a potent inhibitor of BMP signaling (Tuan, 2004). The chondrocytes then begin forming cartilaginous tissue matrices that will be the model for future bones. After the cartilaginous scaffolds of the bones form, the skeletal structures begin to mature through endochondral ossification, one of two mechanisms

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by which bone forms during development. Proliferating cartilage cells in the center of the bone models begin expressing Indian hedgehog (IHH). IHH, parathyroid hormone-related protein (PTHrP), and multiple FGFs coordinate the proliferation of chondrocytes as well as their further maturation into hypertrophic chondrocytes (St-Jacques et al., 1999; Day and Yang, 2008; Ornitz, 2005; Kronenberg, 2006). As the hypertrophic chondrocytes enlarge, they begin to both degrade and mineralize the surrounding matrix before dying and degenerating. Triggering bone formation is the invasion of blood vessels, which bring along osteoclasts to dispose of disintegrating tissue and osteoprogenitors that differentiate into osteoblasts and begin building the bone (Alberts et al., 2002; Hole and Koos, 1994). This close relationship between angiogenesis and osteogenesis has implicated VEGF as playing a major role in the coordination of endochondral ossification (Dai and Rabie, 2007). Osteoblasts deposit a bed of collagenous osteoid that is subsequently mineralized to mature bony tissue, working from the center of the developing bone forming an ossification front that continues to push the chondrocytes outward (Gilbert, 1997). In long bones, the chondrocytes, begin bulging to form heads, and the progression of the ossification front slows. As blood vessels penetrate the heads, a secondary ossification center forms that again ossifies, pushing the chondrocytes outward. Where the secondary and primary ossification fronts meet remains cartilaginous and forms the epiphyseal growth plate, making it possible for further bone extension, until adulthood when bone growth is completed (Gilbert, 1997). Chondrocytes pushed to the edge of the long bones form the articular cartilage, driven in part by TGFβ-induced ECM formation (Eames et al., 2003), and cells remaining outside the bone differentiate to form the periosteum (Hole and Koos, 1994). The second mechanism of bone development, exemplified in the plates of the skull, is intramembranous ossification. Membrane-like layers of primitive connective tissues first appear at the site where bones are to be grown. Mesenchymal cells derived from the neural crest interact with the extracellular matrix of the head epithelial cells to form the bones (Gilbert, 1997). Signaling during these mesenchymal–epithelial interactions includes the BMPs (Gilbert, 1997), TGFβs (Kanaan and Kanaan, 2006), and high Wnt signaling within the mesenchymal condensates to induce osteoblast differentiation (Day and Yang, 2008). Again, vessel formation is critical, as the mesenchymal cells condense around and begin bone formation immediately adjacent to capillaries (Gilbert, 1997). Coordinating these mesenchymal cells through proliferation to differentiation to osteogenesis are members of the FGF (FGF18 and FGF2) and BMP families. Bone formation from differentiating osteoblasts commences the same as in endochondral ossification, forming bony islands that eventually connect to form the plates of the skull.

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During skeletal development, it is interesting that separate embryonic lineages differentiate to form the long bones and craniofacial bones. The long bones originate from the lateral plate mesoderm of the developing embryo, giving these cells a mesenchymal lineage, while craniofacial bones are generated from neural crest cells of ectodermal lineage (Alberts et al., 2002). Although not determined, the different heritage of these cells may influence their response to various growth factor stimuli in adult life, and therefore location must be considered during design and testing of implants. As mentioned above, bone growth occurs post-natally at epiphyseal growth plates. Growth at the plates resembles that of endochondral ossification. Proliferating chondrocytes extend the bone at the leading edge of an ossification front, leaving behind hypertrophic chondrocytes forming calcified cartilage. The calcified cartilage is removed by invading osteoclasts and replaced with mature bone by osteoblasts (Gilbert, 1997). Critical in stimulating the growth of the epiphyseal growth plates are factors including growth hormone (GH) and IGF1 (Gilbert, 1997). As during development, IHH and PTHrP constitute a feedback loop regulating chondrocyte proliferation and differentiation (van der Eerden et al., 2003). Multiple BMPs, including BMP2, BMP4, and BMP7, also play a role, as well as members of the FGF family (van der Eerden et al., 2003). Although a daunting list, it is important to be aware of the multitude of cytokines and the processes they coordinate during development. As the molecular mechanisms underlying bone repair after injury are elucidated, the number of similarities with embryonic developmental processes continues to compound (Gerstenfeld et al., 2003; Dimitriou et al., 2005).

2.4

Bone homeostasis

The structure of bone tissue at any given anatomical location is no mistake, and reflects the optimal mass and morphology to develop the strength required, as well as the optimal shape to satisfy its local load-bearing requirements (Guyton and Hall, 1996; Frost, 1987). These changes in structure, as well as repair of bone microdamage, are accomplished through highly coordinated processes involving the osteoblasts and osteoclasts, and disruptions in this delicate balance is attributed to many bone diseases (Marks and Odgren, 2002). This remodeling occurs via cooperation of many cell types in a basic multicellular unit (BMU), morphologically resembling a cutting cone – osteoclasts leading the way in resorbing bone, osteoblasts following laying down new bone, and blood vessels at the end providing essential nutrients to the newly formed bone (Jilka, 2003). The exact cellular and molecular signals that initiate bone remodeling are unknown, but either resident osteocytes or bone lining cells are likely responsible (Jilka, 2003).

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The signaling cells first recruit osteoclast progenitors to the site of remodeling by releasing chemokines, including monocyte chemoattractant protein-1 (MCP-1, also known as CCL2) and stromal cell-derived factor (SDF-1, also known as CXCL12) (Matsuo and Irie, 2008). Initial differentiation of osteoclasts is induced by the same osteocytes or stromal cells by cytokines, including macrophage-colony stimulating factor (M-CSF, also known as CSF-1) and receptor activator for nuclear factor κB ligand (RANKL) that bind to their correlative receptors on osteoclast progenitors (Jilka, 2003; Matsuo and Irie, 2008). Allowing for fine control of the process, signaling cells also release osteoprotegerin (OPG), which binds to RANKL and prevents its signaling to osteoclast progenitors (Jilka, 2003). Systemic factors such as parathyroid hormone (PTH) and 1,25-vitamin D3 as well as paracrine factors such as interleukin (IL)-1, IL-6, IL-11, and tumor necrosis factor (TNF) can also influence the extent of osteoclast differentiation. Osteoclasts then initiate the process of ‘coupling’ with osteoblasts to form the BMU, through direct cytokine release, cell-to-cell signaling, or liberation of cytokines from resorbed bone matrix (Matsuo and Irie, 2008). Critical in this process are the familiar skeletal cytokines including BMPs, IGFs, and TGFβ (Jilka, 2003; Matsuo and Irie, 2008), as well as a few other candidates including PDGF, hepatocyte growth factor (HGF), and Wnts (Matsuo and Irie, 2008). TGFβ1 is important in the chemotaxis and proliferation of progenitor cells as well as initial extracellular matrix deposition (Dimitriou et al., 2005), but seems to exert inhibitory effects on final cell differentiation and maturation (Alliston and Derynck, 2000). Reinforcing the idea that proper temporal expression of cytokines is critical, TGFβ1 promotes the production of critical initial extracellular matrix components, such as collagen Type I, osteopontin, and osteonectin (Alliston and Derynck, 2000; Lu et al., 2001), while possibly inhibiting some of the late stage markers of osteoblast differentiation. BMPs seem to induce differentiation of immature mesenchymal cells (Lou, 2001), but the BMPs, especially the popular BMP-2, have been shown to exhibit chemotactic and proliferative effects as well (Dimitriou et al., 2005; Lind, 1998). The IGFs work throughout the bone remodeling process, acting potently as survival factors to prevent apoptosis of differentiating and mature osteoblasts (Jilka, 2003). When the osteoblasts complete bone formation and the new bone is properly vascularized, the bone remodeling process is terminated by either apoptosis of the osteoblasts or their terminal differentiation to osteocytes or bone lining cells (Matsuo and Irie, 2008). It is interesting to note the two completely different lineages from which osteoblasts and osteoclasts arise. The mononucleated osteoblasts originate from progenitor cells, likely located in the bone marrow or periosteum (Marks and Odgren, 2002; Gerstenfeld et al., 2003; Caplan, 1991b; Pittenger et al., 1999; Alhadlaq and Mao, 2004). Osteoclasts originate from the same

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precursors as blood cells, and are formed by fusion of these mononuclear precursors (Gilbert, 1997; Jilka, 2003). The involvement of such seemingly unrelated lineages (blood vs. bone) exemplifies how complex and orchestrated bone homeostasis is.

2.5

Bone repair after injury

The events of bone repair after injury (Fig. 2.1) as alluded to above share multiple cellular and molecular events with development and homeostasis, and the process has been referred to as a regeneration that progresses like a post-natal developmental process (Gerstenfeld et al., 2003; Dimitriou et al., 2005). Multiple tissues are required for successful repair and progress in overlapping phases including the inflammatory, stabilization, repair, and remodeling phases (Fig. 2.1). Although there is evidence for key involvement of other lineages, we will focus on the cellular response of the bone cells after injury, which include five key cellular steps that must occur for rapid, successful regeneration of bone tissue: cell recruitment and chemotaxis to injury site, cell proliferation, extracellular matrix deposition, cell differentiation, and mineralization and maturation (Puleo and Nanci, 1999; Greenhalgh, 1996; Dimitriou et al., 2005). The inflammatory phase is the first phase of bone repair after injury and initiates many critical processes to set in motion the complex events to follow (LeGeros and Craig, 1993; Puleo and Nanci, 1999; Kienapfel et al., 1999; Probst and Spiegel, 1997), as demonstrated by the ability of antiinflammatory drugs to significantly decrease bone ingrowth in orthopedic implants (Cook et al., 1995; Goodman et al., 2002). The disruption of many blood vessels results in the activation and aggregation of platelets, leading to the formation of a blood clot, called a hematoma, which is the initial stabilization of the site and repository of many cytokines (Probst and Spiegel, 1997). White blood cells, leukocytes and neutrophils, soon reach the injury site, along with monocytic phagocytes, which differentiate into macrophages. Aside from fighting off infection, these macrophages act as growth factor factories, manufacturing and releasing approximately 100 biologically active substances (Probst and Spiegel, 1997). Apart from the immune phenotypes, disruption of the bone matrix itself leads to release of cytokines, such as TGFβ1 that exists in a latent form in high quantities in bone until released after injury (Roberts, 2000). Taken together, a plethora of coordinating cytokines and growth factors are present at the initial wound site in large quantities, among them the skeletal cytokines PDGF, BMPs, TGFβs, FGFs, and IGFs (Probst and Spiegel, 1997). TGFβ1 seems especially important, as it is synthesized and released in high amounts by the immune cells, is released from the disrupted bone matrix, and induces positive regulation by resident osteoblasts and

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Progenitor Oc

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Recruitment/ chemotaxis Proliferation ECM deposition Differentiation Mineralization/maturation

2.1 Bone repair after titanium implant placement. Initial placement of a titanium implant initiates a tissue cascade that includes the inflammatory phase, repair phase, and remodeling phase. After the inflammatory phase is nearly complete, these tissue responses are regulated largely by bone cells. Recruitment and chemotaxis of progenitor cells from various areas of bone, including periosteum, endosteum, and bone marrow, to the injury site is the first cellular phase. The cells then begin to proliferate and deposit extracellular matrix (ECM) to stabilize the injury site. Upon stabilization and blood vessel ingrowth, the progenitor cells differentiate to osteoblasts (Ob) and begin to lay down immature, unorganized woven bone. During the final remodeling phase, osteoclast (Oc) progenitors are recruited that subsequently differentiate and begin to digest the immature bone. Osteoblasts (Ob) follow closely behind, depositing mature, organized lamellar bone. Some osteoblasts that become entrapped in bone matrix differentiate to osteocytes (Oy). Various growth factors, including TGFβs, BMPs, and IGFs, coordinate this process and can exert different spatial and temporal effects.

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osteoprogenitors (Probst and Spiegel, 1997; Linkhart et al., 1996; Roberts, 2000; Alliston and Derynck, 2000). Once the initial inflammatory phase is completed, stabilization of the injury site continues in the proliferative phase. In fractures, granulation tissue is often first laid down, then replaced by cartilage, and finally is replaced by bone (Probst and Spiegel, 1997). The formation of connective tissue is accomplished by the migration, proliferation, extracellular matrix deposition, and differentiation of progenitor cells (Davies, 2003; Caplan, 1991a). The origin of these progenitors is currently unknown, but many cellular pockets are culprit (Dimitriou et al., 2005) – multipotent mesenchymal stem cells found in bone marrow (Caplan, 1991b; Pittenger et al., 1999; Alhadlaq and Mao, 2004), osteoprogenitor cells of the periosteum and endosteum covering the outer and inner surfaces of bone (Hutmacher and Sittinger, 2003; Gerstenfeld et al., 2003), bone lining cells on the surfaces of bone and in bone pockets (Marks and Odgren, 2002), or multipotent pericyte cells surrounding blood vessels (Collett and Canfield, 2005; Doherty and Canfield, 1999). The periosteum seems to be the major provider of cells for the external callus, and multipotent mesenchymal stem cells of the marrow appear to be the primary cell source for the internal callus (Gerstenfeld et al., 2003; Probst and Spiegel, 1997; Hole and Koos, 1994). Besides providing cellular help, the periosteum cells release cytokines that stimulate differentiation of multipotent cells to osteoblastic phenotypes such as BMP-2, as well as driving the ingrowth of new blood vessels by VEGF and PDGF (Gerstenfeld et al., 2003). After the initial inflammation phase, the levels of TGFβ1, among other factors, remain elevated (Gerstenfeld et al., 2003), suggesting continued and varied roles throughout the fracture healing process. Of critical importance in the proliferative phase is the growth of new blood vessels, angiogenesis, at the wound site. Bone can form only where there is an adequate blood supply available (Probst and Spiegel, 1997). Angiogenesis also depends on the many cytokines and other biologically active substances released by inflammatory cells. Peptides such as TGFβ, FGF2, PDGF, and VEGF have all been shown to play a part in stimulating angiogenesis (Probst and Spiegel, 1997). The mineralization and maturation phase is the final step in bone repair. After suitable mechanical stabilization is achieved from extracellular matrix deposition, osteoprogenitor cells differentiate into osteoblasts and form a bone matrix at the injury site. Differentiation factors such as TGFβ and BMPs are elevated and critical extracellular matrix components such as osteocalcin and osteopontin begin to be expressed. Much of this bone formation is uncoordinated, leading to the formation of mainly woven bone at the injury site. Transition from woven to organized lamellar bone follows the pattern of homeostatic bone remodeling, and results in bone

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strong enough to restore biomechanical competence (Probst and Spiegel, 1997).

2.6

Bone and joint disease

The skeletal system is subject to many disorders that diminish its function, such as trauma, chronic diseases, tumor resection, or congenital anomalies. These disorders compromise and, if left untreated, usually incapacitate the multiple functions of the skeleton, exhibited clinically as increased fractures or debilitating pain. Musculoskeletal injury affects more than 28 million patients in the United States every year and has a $254 billion impact on the US economy per year (USBJD, 2006). Bone homeostasis is a delicate balance between bone formation and resorption, and if disrupted can lead to many disease states (Marks and Odgren, 2002). Osteoporosis is clinically defined as a symptomatic, generalized decrease in bone mass (McCarthy and Frassica, 1998). Osteoporosis results in increased fracture risk, mostly in the vertebrae and hip (McCarthy and Frassica, 1998), and because of impaired bone repair in these patients, fractures often heal slowly, if at all. Like almost every other tissue of the body, bone is subject to cancerous growth of its cells, although to a much less degree than other tissues, with malignant bone tumors being rare (McCarthy and Frassica, 1998). Osteosarcomas are the most common primary malignant bone tumor, usually affecting children and young adults (McCarthy and Frassica, 1998). Luckily, osteosarcomas can usually be reliably removed from the body with a longterm survival rate afterwards. However, complete resection can sometimes require removal of large portions of bone that cannot heal on its own. In human joints, degenerative diseases can progress to severe states leading to pain and debilitation, necessitating treatment (Hayes et al., 2001; Buckwalter, 2002; Gay et al., 2002; Gravallese, 2002). Osteoarthritis, a degenerative disease characterized by loss of articular cartilage (McCarthy and Frassica, 1998), and rheumatoid arthritis, an inflammatory autoimmune disease that destroys a patient’s own cartilage, are both joint diseases that usually lead to debilitating pain or physical disability (McCarthy and Frassica, 1998). Although normally resulting from problems with the articular cartilage, treatment of these diseases often requires total joint replacements that engage bone sites for strength and stability.

2.7

Current treatment options and total joint replacements

Diagnosed and treated early, bone and joint diseases often can be controlled or even reversed with non-invasive methods such as drug or hormone

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treatments (McCarthy and Frassica, 1998). Unfortunately, these diseases often progress to a late stage, characterized by severe pain and debilitation that inhibits or severely limits normal daily activity, requiring more radical and invasive strategies to restore an acceptable quality of life. At this point of disease progression, current medical approaches dictate full replacement. Aside from total joint replacements, load-bearing plates are often used to stabilize bone after fracture or tumor resection. In comparison with donor site morbidity and pain in association with autologous tissue grafting, synthetic materials have the advantage of ready and endless supply without any sacrifice of donor tissue. Therefore, replacement using synthetic biomaterials such as metals or ceramics has become the preferred treatment for total joint replacement. The current ‘gold standard’ in metals for treatment of skeletal disorders is the titanium implant. Titanium combines high strength and excellent biocompatibility (Long and Rack, 1998; Ratner et al., 1996), arising from a highly inert passivating layer of titanium oxide (TiO2) that forms instantly on exposure to air (Ratner et al., 1996). This passivating layer also is responsible for titanium implants’ remarkable capacity for integration with host bone (Albrektsson et al., 1981; Kienapfel et al., 1999), which occurs first by bone ingrowth in which bone forms within the irregular surfaces of the titanium (Kienapfel et al., 1999) and subsequent osseointegration, defined as ‘direct contact between living bone and an implant on the light microscope level’ (Brånemark et al., 1969). Titanium has high success rates of initial anchorage (over 90%) and has been shown to co-exist with host bone for the life of the patient (Ashley et al., 2003). After insertion, the host bone response to the titanium implant is paramount for a successful outcome (Misch, 1993). Implant integration proceeds almost identically to that of bone repair (Fig. 2.1), practically down to the molecular level (Kojima et al., 2008; De Ranieri et al., 2005a; Schierano et al., 2005). Much like fracture repair, titanium implants can integrate to form a strong and stable interlock with host bone. However, unlike fracture repair, a high degree of initial fixation at surgery is critical for successful performance of the implanted device (Puleo and Nanci, 1999), as any relative motion between implant and bone can result in permanent fibrous encapsulation and failure of the implant (Szmukler-Moncler et al., 1998; Brunski, 1999).

2.8

Current challenges of titanium implants

Titanium implants predictably restore mechanical function, yet failures still remain in the area of initial and long-term integration of the implant into host bone (Merickse-Stern et al., 2001; Greenfield et al., 2002). Inadequate

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bone ingrowth and failure to maintain adequate peri-implant bone density because of osteolysis at the bone–implant interface still lead to many failures of titanium implants, requiring painful revision surgeries (Saleh et al., 2004). Although the precise mechanisms of implant failure are not well understood, the ultimate direct cause is either initial or long-term failure of implant integration with host bone, resulting in the implant loosening and ultimately secondary surgery. Aseptic loosening of the implant is the most common cause for long-term revision surgery (McCarthy and Frassica, 1998). Loosening occurs through multiple mechanisms. Like any artificial biomaterial, the implant is subject to wear and tear from the constant mechanical loading that is applied to it, leading to a gradual loosening (McCarthy and Frassica, 1998). A biologic response occurs via one of two mechanisms. During loading, particles of varying sizes are released both from the titanium implant itself and from the plastic (polyethylene) acetabular cup that usually accompanies the implant. This particle wear from polyethylene is taken up by both macrophages for small particles and giant cells for the larger particles, which invokes macrophages to release osteoclast-specific cytokines, inducing osteolysis at the bone–implant interface (McCarthy and Frassica, 1998; McKoy et al., 2000; Ingham and Fisher, 2000). Secondly, the mis-match of biomechanical properties between titanium and bone (the modulus of elasticity of titanium is about ten times greater than that of cortical bone and one hundred times greater than that of cancellous (An, 2000; Ratner et al., 1996)) can lead to abnormally low stress transfer to surrounding bone, initiating a remodeling response from osteoclasts and bone disuse atrophy (McKoy et al., 2000). This condition is known as stress shielding (McKoy et al., 2000; Sumner and Galante, 1992). The most important factors in the relief of stress shielding are the design of the implant to distribute load equally to surrounding bone and the modulus of elasticity of the material used in the implant (McKoy et al., 2000). Because of these long-term issues with titanium implants, it may seem that failure of a titanium implant for joint replacement is inevitable. However, these processes are gradual, occurring over many years (McCarthy and Frassica, 1998). A strong initial bone formation response providing good fixation of the implant, combined with long-term delivery of osteogenic stimuli, such as cytokines, to counteract the bone resorption response could significantly extend the life of titanium implants. Therefore, modulating the complex tissue and cellular bone repair responses at the bone–implant interface to improve the implant fixation process is desirable to improve both short- and long-term efficacy of titanium implants.

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2.9

Current titanium modifications for improved integration

Modification of titanium implants to improve the integration process is by no means a new idea, and much research using various approaches has been reported. The performance of an implant depends on two major factors: the behavior of the material in the host and the response of the host to the implant (Puleo and Nanci, 1999). The behavior of the material in the host depends on the biocompatibility of material implanted. In the case of titanium, once exposed to oxygen in ambient air or in the body, titanium forms a passivating layer of titanium oxide (TiO2) almost instantly, making the surface bioinert and allowing bone ingrowth (Ratner et al., 1996). Different treatments of the surface (Sul et al., 2001; Pan et al., 1998) and even different methods of sterilization (Binon et al., 1992) can affect the thickness of this oxide layer, which in turn affects cellular response to the titanium implants (Pan et al., 1998). Tailoring the response of the host to the implant has focused mainly on surface modification of implanted titanium (Puleo and Nanci, 1999). Perhaps the simplest and most successful methods have been morphological alterations of surface morphology and roughness of the implant. Surface roughening to create a porous titanium surface on the macro- and micro-scales has been accomplished by many methods, including machine-smoothing, grit-blasting, sand-blasting, acid etching, and plasma-spraying (Boyan et al., 1996; Pilliar, 2005). Use of these surface roughened implants has become commonplace, and the idea of ‘bone ingrowth’ into the porous spaces of titanium implant surfaces has become an important first step in the integration of titanium implants with host bone (Kienapfel et al., 1999; Pilliar, 2005). Varying degrees of surface roughness have been shown to effect cellular response to titanium implants (Jayaraman et al., 2004; Brett et al., 2004; Mante et al., 2003; Sittig et al., 1999; Schwartz et al., 1997; Boyan et al., 1996) and ultimately bone ingrowth (Boyan et al., 1996) in both dentistry and orthopedics (Puleo and Nanci, 1999; Kienapfel et al., 1999), with increasing surface roughness correlating to increased cell proliferation (Mante et al., 2003) and bony ingrowth (Frenkel et al., 2002; Groessner-Schreiber and Tuan, 1992; Wong et al., 1995; Thomas and Cook, 1985) up to an optimal value (Ronold et al., 2003). On the other hand, macroscopic or microscopic roughening of the titanium allows development of stress and strain concentrations, as is the case for some macroscopic screw threads (Van Oosterwyck et al., 1998), and if these grow too large, bone ingrowth can be impeded (Qin et al., 1996). Physiochemical methods of surface modification refer to increasing surface free energy in order to increase tissue adhesion; methods used

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include glow discharge and charging the surface of the implant (Puleo and Nanci, 1999). An important approach in this category is coating implants with osteoinductive and osteoconductive materials, the most established being calcium phosphate derivatives (Lavelle et al., 1981; Kay, 1993) such as hydroxyapatite (HA) and tricalcium phosphate (TCP). These coatings have been prepared using various techniques, including precipitation, plasma spraying, and sputtering (Hulshoff et al., 1996; Jansen et al., 1993; Wen et al., 1998), and regardless of coating technique have consistently shown increased implant fixation when compared to non-coated titanium (Soballe et al., 1992; Maxian et al., 1993; Tisdel et al., 1994; Dean et al., 1995). In vitro assays have shown improved mineralization (Hulshoff et al., 1996; Morgan et al., 1996; ter Brugge et al., 2002) and increased osteoprogenitor differentiation (ter Brugge et al., 2002; Hulshoff et al., 1996) of calcium phosphate surfaces, with the degree of crystallinity of calcium phosphate playing a major role in cell response (Hulshoff et al., 1996; Berube et al., 2005). However, calcium phosphate coatings are not perfect, in that the mechanical strength between the calcium phosphate and titanium is often of questionable and variable strength (Filiaggi et al., 1991; Ergun et al., 2003) and the calcium phosphate may also stimulate osteoclast mediated bone resorption (Gottlander et al., 1997). Many newer approaches of surface modifications fall under the category of biochemical methods, in which proteins, enzymes, or peptides are immobilized on biomaterials for the purpose of inducing specific cell and tissue responses (Puleo and Nanci, 1999). Coatings in this category include collagen (Kim et al., 2005; Park et al., 2005), the RGD peptide (Ferris et al., 1999; Huang et al., 2003), and fibronectin (Yang et al., 2003), and have resulted in improvements for both in vitro assays and short-term in vivo implantations (Ferris et al., 1999; Schliephake et al., 2005; Morra et al., 2005; Rammelt et al., 2004). These surface modifications show promise for improved implant fixation, although long-term and human studies need still to be performed. During past decades, the premise of the design of synthetic tissue implants has been to use inert and bulk materials that permit the integration of host tissue. This approach allows titanium implants to retain their high strength; however, this high strength diverts functioning mechanical stress necessary for healthy peri-implant bone and leads to stress shielding related osteoclastogenesis and osteolysis (McKoy et al., 2000; Sumner and Galante, 1992). Mechanics of materials theory dictates that the removal of material from the bulk decreases the apparent modulus of elasticity (Hibbeler, 1997), reducing the disparity of the bone to implant modulus mis-match (Thelen et al., 2004). Removal of pieces of the bulk of titanium implants would also potentially lead to a greater degree of interlock between the implant and host, presumably leading to better implant fixation.

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Because of the aforementioned reasons, the use of a hollow implant design seems advantageous. Hollow designs such as meshes and cages have been tested for their integrative potential, but largely in non loadbearing applications (Vehof et al., 2002; Aytac et al., 2005). Load-bearing hollow implants in dentistry have existed almost as long as solid implants (Schroeder et al., 1976; Merickse-Stern et al., 2001; Misch, 1993). Although mechanical failures of these implants have been reported (Schwarz, 2000; Levine et al., 1999), the implants have performed at least as well as other implant designs (Buser et al., 1997). New tissue engineering technologies in the areas of drug and cell delivery have miniaturized delivery vehicles to the point where the hollow core of a titanium implant could also be used as a storage space for osteogenic drugs and/or cells.

2.10

Mimicking nature toward achieving titanium ‘biointegration’: cytokines and implants

With further biological and molecular understanding of bone development, growth, and repair, the importance of soluble cytokines has been realized in orthopedics. Identification of cytokines that regulate the key steps of titanium implant integration (cell recruitment and chemotaxis to injury site, cell proliferation, extracellular matrix deposition, cell differentiation, and mineralization and maturation) (Alliston and Derynck, 2000; Probst and Spiegel, 1997; Schwartz et al., 1997; Puleo and Nanci, 1999; Steinbrech et al., 2000; Hunziker and Rosenberg, 1996) has opened the door for powerful molecular based approaches that mimic the natural healing process. Although not a defined cocktail, platelet-rich plasma contains a multitude of cytokines expressed during the inflammation phase, and has been used to increase the volume of peri-implant bone in rat tibiae (Fontana et al., 2004) and increase bone ingrowth when delivered to peri-implant defects via a demineralized freeze-dried bone carrier (Sanchez et al., 2005). Injection of FGF2 increased new bone formation and connection to titanium implants implanted in rat medullary cavities (Takechi et al., 2008). Osteogenic factors such as bone morphogenetic proteins (BMPs), transforming growth factor β (TGFβ), and platelet-derived growth factor (PDGF) have repeatedly led to improvements in many bone-implant integration parameters (Wikesjo et al., 2005; Soballe et al., 2004; Lynch et al., 1991; De Ranieri et al., 2005b; Jansen et al., 2005). Use of TGFβ1 has been shown to improve bone repair and titanium implant healing (Sumner et al., 1995; Hong et al., 2000a,b; Yamamoto et al., 2000; Jansen et al., 2005; Ehrhart et al., 2005). In applications where there are existing gaps, such as augmentation of bone at implant sites or peri-implant defects, increased regeneration of bone, often at comparable levels to bone grafts, has been accomplished with delivery of TGFβ and BMP-2 via various scaffolds (Cochran et al., 1999;

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Clokie and Bell, 2003; Clarke et al., 2004; Yamamoto et al., 2000; Vehof et al., 2002; Jansen et al., 2005). Adsorption of TGFβ and/or BMP to tricalcium phosphate or hydroxyapatite coatings on titanium implants leads to enhancement of bone ingrowth in gap models (Sumner et al., 1995, 2006; Lind, 1998; Zhang et al., 2004). Obviously, cytokines possess great potential to improve titanium–bone integration.

2.11

Growth factor delivery: why is controlled and sustained release important?

As discussed, the interplay of soluble cytokines is an elegant process, both spatially and temporally. One may question why improved delivery systems are required for cytokines when such positive outcomes as those above are achieved through simple injection to the implant site or adsorption to the titanium implant. Unfortunately, several limitations and areas for improvements are evident in these systems. First, cytokines traditionally exhibit a short half-life in vivo before diffusion from the injury site or denature from enzymes (Moioli et al., 2007; Alliston and Derynck, 2000). Therefore, large amounts of TGFβ are required for improvement of key bone ingrowth parameters (De Ranieri et al., 2005b; Jansen et al., 2005; Ehrhart et al., 2005; Sumner et al., 1995). From an application perspective, the high expense of most cytokines makes this a major drawback for widespread use. More importantly, large amounts of delivered skeletal growth factor decreases safety and efficacy, potentially leading to side effects such as ectopic bone formation (Wildemann et al., 2004) or oncogenic transformation. Second, mimicking nature through controlled temporal and spatial release will most likely produce more optimal and efficacious effects. Injected or adsorbed proteins are rapidly exhausted whereas levels of many cytokines, including TGFβ and BMP2, remain elevated throughout the fracture repair process (Gerstenfeld et al., 2003). Designed delivery would take advantage of individual cytokines for optimal use; for example, early delivery of TGFβ for mobilization and expansion of osteoprogenitors followed by BMP2 for maturation and mineralization of the resulting pool. Fortunately, biomaterials research has provided a multitude of growth factor delivery systems to overcome these challenges, focused mainly on the use of biodegradable polymers. Many of the most common systems incorporate poly-L-lactic acid (PLLA), poly-glycolic acid (PGA), or the co-polymer poly-lactic-co-glycolic acid (PLGA) (Moioli et al., 2006; Cohen et al., 1991; Lu et al., 2000; Wildemann et al., 2004). As these polymers hydrolyze in the in vivo environment to their biocompatible byproducts, encapsulated cytokines are slowly released to surrounding tissue. However, as

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plastics, these systems lack the required mechanical strength for functional loading of the skeletal system (Ratner et al., 1996). Using special techniques, multiple polymers have been prepared as microparticles, or microspheres, that can encapsulate water soluble proteins (Cohen et al., 1991; Iwata and McGinity, 1992) including cytokines (Lu et al., 2001; Oldham et al., 2000; Moioli et al., 2006; Clark et al., 2008). The small size of polymer microspheres allows for their incorporation into stronger materials, such as titanium, while still exhibiting slow release over an extended period of time. Injection of these microparticles has also been shown to be feasible (Jiang et al., 2005), which could be a potential route to deliver osteogenic cytokines if maintenance of implant fixation was required in the long-term. Using these approaches, we have fabricated titanium implants with a hollow core and interspersed macropores; these hollow implants maintain the biomechanical strength required for load-bearing applications but provide space for tissue engineered scaffolds delivering bioactive molecules or cells (Fig. 2.2a) (Clark et al., 2008). For incorporation into the implant, PLGA microspheres (MS) averaging 64 μm in diameter (Fig. 2.2b) were fabricated to release approximately 100 ng TGFβ1 in a controlled fashion up to four weeks (Fig. 2.2c). Dosage was calculated from release kinetics and corresponded to 0.005 μg/mm3 of injury volume, equivalent to that found using controlled release of TGFβ1 to repair skull defects (Ueda et al., 2002; Yamamoto et al., 2000). After injection into a gelatin carrier, the PLGA microspheres could be subsequently packed into the hollow core of the titanium implant. The hollow implants were placed unicortically into the humeri of adult New Zealand white rabbits (Fig. 2.2d) and allowed to integrate for four weeks. As controls, placebo microspheres (no TGFβ1) as well as 1 μg and 100 ng adsorbed to gelatin carrier (rapid release) were used. After analysis using scanning electron microscopy (SEM), histology, image-assisted histomorphometry, and microcomputed tomography (μCT) imaging, it was demonstrated that incorporation of TGFβ1 in the hollow implant led to significantly increased osseointegration parameters (Figs 2.2e–l). Increased bone apposition, measured as bone-to-implant contact (BIC), to the titanium surface was observed in the 1 μg TGFβ1 gelatin and 100 ng TGFβ1 controlled release MS groups, as compared to 100 ng gelatin and placebo MS groups (Figs 2.2e–h, arrows). Concurrently, increased woven bone (WB) density, measured as bone volume/tissue volume (BV/TV), was seen within the macropores of the 1 μg TGFβ1 gelatin and 100 ng TGFβ1 controlled release MS groups, as compared to 100 ng gelatin and placebo MS groups (Figs 2.2i–l). The measured increases of 96% for BIC and 50% for BV/TV via control-released TGFβ1 over placebo MPs were comparable to bone ingrowth studies using growth factor adsorption, with the important difference that the control-release approach reduced the required drug

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Ti 500 μm

2.2 Hollow titanium implants with a hollow core and macropores were custom fabricated (a). Poly-lactic-co-glycolic acid (PLGA) microspheres encapsulating TGFβ1 with an average diameter of 64 μm were constructed (b) that demonstrated release kinetics of a quick initial burst at three days and sustained release up to four weeks (c). To compare the outcome to a rapid release system, 1 μg or 100 ng of TGFβ1 was also adsorbed to a gelatin sponge carrier (f, g, j, k). After four weeks of implantation in rabbit humeri (d), controlled delivery of TGF-β1 led to increases in bone-to-implant contact (BIC) (e–h, arrows) and woven bone (WB) volume within macropores (BV/TV) (i–l) as compared to placebo control spheres (e and i). Approximately ten times more TGFβ1 was required for rapid release from the gelatin sponge (g and k) to obtain results comparable with controlled release using PLGA microspheres (h and l). Scanning electron microscopy (SEM) imaging. Ti, titanium implant; MS, microspheres.

dose by ten-fold (Sumner et al., 1995; Ehrhart et al., 2005; Schmidmaier et al., 2001). These results could have significant implications in potential reductions of cost and toxicity of in vivo delivered biological cues, and established a proof-of-concept for porous implants as delivery vehicles for controlled release of microencapsulated bioactive cues (Clark et al., 2008).

2.12

Future trends

Growth factor delivery for enhancement of orthopedic implant integration is a field just beginning to emerge, part of the larger fields of tissue engineering and regenerative medicine incorporating materials science, cell and

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molecular biology, biochemistry, and bioengineering. Success in enhancing orthopedic implant integration and efficacy requires a concerted effort from all these areas in refining and improving the above discussed approaches. Biomaterials delivery systems are becoming more and more elegant every day. Poly-lactic-co-glycolic acid (PLGA) microspheres represent a standard in growth factor delivery. Depending on the formulation of PLGA microspheres, especially the ratio of PLA to PGA, controlled release at various rates from weeks to months can be achieved (Moioli et al., 2006; Cohen et al., 1991), and the polymer has already been approved as safe and efficacious for human use by the United States Federal Drug Administration (FDA). Further refinements to PLGA as well as other materials for drug delivery are being actively investigated, including but not limited to poly-ethylene glycol (PEG) derivatives (Stosich et al., 2007; Barralet et al., 2005), oligo(poly(ethylene glycol) fumarate) (Temenoff et al., 2004; Holland et al., 2005), chitosan microspheres (Cho et al., 2004), and hyaluronic acid (Kim and Valentini, 2002; Angele et al., 1999; Bulpitt and Aeschlimann, 1999). More advanced load-bearing biomaterials are also becoming increasingly available. New processing procedures for titanium have led to the advent of microporous titanium and titanium foams (Thelen et al., 2004; Wen et al., 2002), which have porosities conducive for bone formation while still retaining weight-bearing strength. Similar processes exist for ceramics as well (Kim et al., 2004; Baran et al., 2004). Materials such as these may provide hollow implants with more stable biomechanical loading properties, while still maintaining porosity for the delivery of bone enhancing cytokines. Numerous cytokines have been implicated at different spatial and temporal locations during bone growth and repair after injury, including the popular skeletal factors (TGFβs, FGFs, BMPs) (Gerstenfeld et al., 2003; Dimitriou et al., 2005; Kuroda et al., 2005) as well as new candidates (Wnts and Hedgehog) (Day and Yang, 2008; St-Jacques et al., 1999; Franceschi, 2005; van der Eerden et al., 2003). It is unlikely that any single growth factor is optimal in inducing bone repair and integration after titanium implantation. Adaptation of the PLGA microsphere system or use of other systems are beginning to be used for multiple growth factor delivery (Richardson et al., 2001; Simmons et al., 2004; Sumner et al., 2006; Holland et al., 2007). Not to be forgotten, vascularization is critical for successful bone repair, and angiogenic strategies are also being developed to be used alongside osteogenic ones (Richardson et al., 2001; Patel et al., 2008). Incorporation of various cell types alongside medical implants is quickly emerging as a viable treatment option. Osteoprogenitors or mature osteoblasts have been demonstrated to accelerate and enhance the osseointegration of titanium implants (Frosch et al., 2003). The discovery of mesenchymal stem cells capable of multipotential differentiation has been

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a major breakthrough for regenerative medicine, providing a source of highly expandable cells for cartilage, bone, and adipogenic applications (Caplan, 1991a; Alhadlaq et al., 2004; Williams et al., 2003). Mesenchymal stem cells (MSCs) could easily be incorporated into hollow implants alongside cytokines, possibly leading to inside-out bone formation by delivered MSCs and outside-in bone formation by MSCs recruited by delivered cytokines from the external environment (Clark et al., 2008). Although there is still much to be learned on how to control their osteoblastic differentiation, human embryonic stem cells (hESCs) (Thomson et al., 1998; Sottile et al., 2003) and the recently discovered induced pluripotent stem (iPS) cells (Yu et al., 2007; Takahashi et al., 2007) also represent candidates for bone regeneration applications.

2.13

Acknowledgements

This work is supported by NIH grants DE15391 and EB02332 to J.J.M. P.A.C is supported by a post-doctoral fellowship through the UW–Madison Stem Cell Training Program funded by NIH (5T32AGO27566-03).

2.14

Sources of further information and advice

Further information in the fields discussed can be obtained from various locations. Further exploration into biology can be found through various textbooks, such as for developmental biology, Developmental Biology by S. F. Gilbert, 1997, or newer editions (Gilbert, 1997), molecular and cellular biology, Molecular Biology of the Cell by B. Alberts, A. Johnson, J. Lewis, M. Raff, K. Roberts, and P. Walter, 2002, or newer editions (Alberts et al., 2002), bone biology, Principles of Bone Biology by J.P. Bilezikian, L.G. Raisz, and G.A. Rodan, 2002 or newer editions (Bilezikian, 2002), and growth factor biology, Skeletal Growth Factors by E. Canalis, 2000, or newer editions (Canalis, 2000). Further information on biomaterials for medical applications can be found through textbooks such as Biomaterials Science: An Introduction to Materials in Medicine by B. D. Ratner (Ratner et al., 1996) or newer editions. Principles of Tissue Engineering by R. Lanza, R. Langer, and J. P. Vacanti (Lanza et al., 2007) gives a thorough overview of the tissue engineering and regenerative medicine fields. Translational Approaches in Tissue Engineering and Regenerative Medicine by J. J. Mao, G. Vunjak-Novakovic, A. Mikos and A. Atala (Mao et al., 2007) provides a comprehensive coverage of cutting-edge science and technologies regarding implant integration, tissue engineering and translational approaches in regenerative medicine. Many organizations are rich with pertinent information. The National Institutes of Health (NIH, www.nih.gov) and its associated institutes, such

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as the National Institute of Arthritis and Musculoskeletal and Skin Diseases (NIAMS, www.niams.nih.gov), provide background on bone diseases as well as current therapies and approaches. The National Institute of Dental and Craniofacial Research (NIDCR) has a wealth of information regarding dental and craniofacial implants (www.nidcr.nih.gov). Many societies and foundations have been founded to further research in bone diseases and repair that provide background as well as recent advances in various fields – the American Academy of Implant Dentistry (AAID, www.aaid-implant. org), American Academy of Orthopedic Surgeons (AAOS, www.aaos.org), Arthritis Foundation (www.arthritis.org), National Osteoporosis Foundation (NOF, www.nof.org), Orthopedic Research Society (ORS, www.ors. org), Society for Biomaterials (SFB, www.biomaterials.org), and Tissue Engineering and Regenerative Medicine International Society (TERMIS, www.termis.org).

2.15

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mao, j. j., vunjak-novakovic, g., mikos, a. and atala, a. (2007) Translational Approaches in Tissue Engineering and Regenerative Medicine, Boston, Artech House. marie, p. j., debiais, f. and hay, e. (2002) Regulation of human cranial osteoblast phenotype by FGF-2, FGFR-2 and BMP-2 signaling. Histol Histopathol, 17, 877–85. marks, s. c. and odgren, p. r. (2002) Structure and development of the skeleton. In Bilezikian, J. P., Raisz, L. G. and Rodan, G. A. (Eds.) Principles of Bone Biology, Volume 1. 2nd ed. San Diego, Academic Press. matsuo, k. and irie, n. (2008) Osteoclast–osteoblast communication. Arch Biochem Biophys, 473, 201–9. maxian, s. h., zawadsky, j. p. and dunn, m. g. (1993) Mechanical and histological evaluation of amorphous calcium phosphate and poorly crystallized hydroxyapatite coatings on titanium implants. J Biomed Mater Res, 27, 717–28. mcbeath, r., pirone, d. m., nelson, c. m., bhadriraju, k. and chen, c. s. (2004) Cell shape, cytoskeletal tension, and RhoA regulate stem cell lineage commitment. Dev Cell, 6, 483–95. mccarthy, e. f. and frassica, f. j. (1998) Pathology of Bone and Joint Disorders, Philadelphia, W.B. Saunders Company. mckoy, b. e., an, y. h. and friedman, r. j. (2000) Factors affecting the strength of the bone-implant interface. In An, Y. H. and Draughn, R. A. (Eds.) Mechanical Testing of Bone and the Bone–Implant Interface. New York, CRC Press. merickse-stern, r., aerni, d., geering, a. h. and buser, d. (2001) Long-term evaluation of non-submerged hollow cylinder implants. Clinical and radiographic results. Clin Oral Implants Res, 12, 252–9. misch, c. e. (1993) Contemporary Implant Dentistry, Chicago, Mosby. moioli, e. k., clark, p. a., xin, x., lal, s. and mao, j. j. (2007) Matrices and scaffolds for drug delivery in dental, oral and craniofacial tissue engineering. Adv Drug Deliv Rev, 59, 308–24. moioli, e. k., hong, l., guardado, j., clark, p. a. and mao, j. j. (2006) Sustained release of TGFbeta3 from PLGA microspheres and its effect on early osteogenic differentiation of human mesenchymal stem cells. Tissue Eng, 12, 537–46. morgan, j., holtman, k. r., keller, j. c. and stanford, c. m. (1996) In vitro mineralization and implant calcium phosphate–hydroxyapatite crystallinity. Implant Dent, 5, 264–71. morra, m., cassinelli, c., meda, l., fini, m., giavaresi, g. and giardino, r. (2005) Surface analysis and effects on interfacial bone microhardness of collagen-coated titanium implants: A rabbit model. Int J Oral Maxillofac Implants, 20, 23–30. oldham, j. b., lu, l., zhu, x., porter, b. d., hefferan, t. e., larson, d. r., currier, b. l., mikos, a. g. and yaszemski, m. j. (2000) Biological activity of rhBMP-2 released from PLGA microspheres. J Biomech Eng, 122, 289–92. ornitz, d. m. (2005) FGF signaling in the developing endochondral skeleton. Cytokine Growth Factor Rev, 16, 205–13. pan, j., liao, h., leygraf, c., thierry, d. and li, j. (1998) Variation of oxide films on titanium induced by osteoblast-like cell culture and the influence of an H2O2 pretreatment. J Biomed Mater Res, 40, 244–56. park, b. s., heo, s. j., kim, c. s., oh, j. e., kim, j. m., lee, g., park, w. h., chung, c. p. and min, b. m. (2005) Effects of adhesion molecules on the behavior of osteoblast-like

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cells and normal human fibroblasts on different titanium surfaces. J Biomed Mater Res A, 74, 640–51. patel, z. s., young, s., tabata, y., jansen, j. a., wong, m. e. and mikos, a. g. (2008) Dual delivery of an angiogenic and an osteogenic growth factor for bone regeneration in a critical size defect model. Bone, 43, 931–40. pilliar, r. m. (2005) Cementless implant fixation – toward improved reliability. Orthop Clin North Am, 36, 113–19. pittenger, m. f., mackay, a. m., beck, s. c., jaiswal, r. k., douglas, r., mosca, j. d., moorman, m. a., simonetti, d. w., craig, s. and marshak, d. r. (1999) Multilineage potential of adult human mesenchymal stem cells. Science, 284, 143–7. probst, a. and spiegel, h. u. (1997) Cellular mechanisms of bone repair. J Invest Surg, 10, 77–86. puleo, d. a. and nanci, a. (1999) Understanding and controlling the bone–implant interface. Biomaterials, 20, 2311–21. qin, y. x., mcleod, k. j., guilak, f., chiang, f. p. and rubin, c. t. (1996) Correlation of bony ingrowth to the distribution of stress and strain parameters surrounding a porous-coated implant. J Orthop Res, 14, 862–70. rammelt, s., schulze, e., bernhardt, r., hanisch, u., scharnweber, d., worch, h., zwipp, h. and biewener, a. (2004) Coating of titanium implants with Type-I collagen. J Orthop Res, 22, 1025–34. ratner, b. d., hoffman, a. s., schoen, f. j. and lemons, j. e. (1996) Biomaterials Science: An Introduction to Materials in Medicine, New York, Academic Press. richardson, t. p., peters, m. c., ennett, a. b. and mooney, d. J. (2001) Polymeric system for dual growth factor delivery. Nat Biotechnol, 19, 1029–34. roberts, a. b. (2000) Transforming growth factor-beta In Canalis, E. (Ed.) Skeletal Growth Factors. Philadelphia, Lippincott, Williams, and Wilkins. ronold, h. j., lyngstadaas, s. p. and ellingsen, j. e. (2003) Analysing the optimal value for titanium implant roughness in bone attachment using a tensile test. Biomaterials, 24, 4559–64. saleh, k. j., thongtrangan, i. and schwarz, e. m. (2004) Osteolysis: Medical and surgical approaches. Clin Orthop Relat Res, 138–47. sanchez, a. r., sheridan, p. j., eckert, s. e. and weaver, a. l. (2005) Regenerative potential of platelet-rich plasma added to xenogenic bone grafts in peri-implant defects: A histomorphometric analysis in dogs. J Periodontol, 76, 1637–44. scadden, d. t. (2006) The stem-cell niche as an entity of action. Nature, 441, 1075–9. schierano, g., canuto, r. a., navone, r., peirone, b., martinasso, g., pagano, m., maggiora, m., manzella, c., easton, m., davit, a., trombetta, a., amedeo, s., biolatti, b., carossa, s. and preti, g. (2005) Biological factors involved in the osseointegration of oral titanium implants with different surfaces: A pilot study in minipigs. J Periodontol, 76, 1710–20. schliephake, h., scharnweber, d., dard, m., sewing, a., aref, a. and roessler, s. (2005) Functionalization of dental implant surfaces using adhesion molecules. J Biomed Mater Res B Appl Biomater, 73, 88–96. schmidmaier, g., wildemann, b., bail, h., lucke, m., fuchs, t., stemberger, a., flyvbjerg, a., haas, n. p. and raschke, m. (2001) Local application of growth factors (insulin-like growth factor-1 and transforming growth factor-beta1) from a biodegradable poly(D,L-lactide) coating of osteosynthetic implants accelerates fracture healing in rats. Bone, 28, 341–50.

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3 Replacement materials for facial reconstruction at the soft tissue– bone interface E. W E N T R U P-B Y R N E, Queensland University of Technology, Australia; L. G R Ø N DA H L and A. C H A N D L E R-T E M P L E, The University of Queensland, Australia

Abstract: The challenges faced by any tissue repair and regeneration process resulting from either trauma or disease are many and complex. Although it is of course impossible to identify any one anatomical region as being the most demanding in this respect, the craniofacial region surely qualifies. The judicious choice of available, well-defined and tested repair materials to be used in the reconstruction process by the multi-disciplinary team of reconstructive surgeons is critical. This chapter addresses one aspect of facial reconstruction that has been less well addressed in the literature; namely the materials used to repair and regenerate soft tissue both in terms of fillers and in terms of materials used at the hard–soft tissue interface. Key words: soft tissue–bone interface, expanded polytetrafluoroethylene, surface modification, biomineralisation, bioresorbable fillers.

3.1

Introduction

In spite of the beauty and complexity of the magnificently engineered structure that constitutes the human face, it appears that humankind has never been ‘satisfied’ with their faces and how they ‘look’. In his fascinating book, Landau states that documentation dating back to the Minoan Bronze Age around 3500 bc confirms that altering the human head and face occurred in several civilisations and he speculates that more recent evidence dates the practice even further back in the mists of time. The human face is intrinsically linked to a person’s identity. Our fascination with the human face ranges from the artist who continually paints self-portraits – none more famous than Rembrandt who painted over 90 (Osmond, 2000) – to one of the most recognised faces of all, ‘The Mona Lisa’, and the speculation about her identity (Rising, 2008). Scientists such as the renowned Erik Erikson spent a lifetime exploring the meaning of identity; in fact the term ‘identity crisis’ is attributed to him. One aspect of his research involved 51 © Woodhead Publishing Limited, 2010

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Frontal bone

Parietal bone

Sphenoid bone

Nasal bone

Temporal bone Occipital bone

Zygomatic bone Maxilla

Mandible

3.1 Lateral view of skull showing principal bones.

the ‘loss’ of identity in World War II soldiers as a result of facial injury (Landau, 1989). The psychological repercussions and effects on a person’s life of severe deformities resulting from either trauma or genetic malformations cannot be underestimated. As evidenced by the recent global publicity afforded to the first reports of facial or ‘near-total’ facial transplants, interest in restoring the aesthetic appearance is seen as just as important as restoration of function (BBC, 2009; Gonzalez, 2008; AFP, 2008). Figure 3.1 shows some of the principal bones of the craniofacial skeleton which forms the foundations for the aesthetic features of the human face. To quote from a presentation by Fialkov et al. (2000) at the Bone Engineering Workshop, held in Toronto in 1999, ‘The morphology of the entire facial skeleton, in particular the upper region, has significant cultural, sexual, and social implications. Concepts of beauty, youth and intelligence are associated with particular facial morphologies, and vary with ethnicity. Much of this morphology is based on the underlying bony structure of the facial skeleton.’ The main functions of the human craniofacial skeleton are threefold. First, the multilayered bone framework provides protection for vital

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3.2 Anterior cut away view of the face, showing the soft tissue (left), and a radiograph showing the underlying bone structure (right).

structures such as the ocular and aural systems, the central nervous system and the upper aero-digestive tract. Second, it also forms the foundations for the aesthetic features of the human face. The third and final function is that, through providing the structural framework, lever and fulcrum, it renders mastication possible. Another important component of one’s facial features is the soft adipose (subcutaneous fat) tissue that forms the interface between the bones and the skin. Figure 3.2 shows an anterior cut away view of the face showing the soft tissue as well as a radiograph showing the underlying bone structure. The anatomical interrelationship between these ‘structures’ is a complex one and this, of course, means that in order to fulfil their functions, their repair and regeneration are also interdependent. Plastic surgery constitutes an enormous and sometimes much criticised industry, but it also includes the extremely important challenge of trauma repair. One example where facial surgery may be required and which is particularly relevant to this chapter is the case of human immunodeficiency virus (HIV)-related lipoatrophy (LA). As far as can be ascertained, Carr et al. (1998) were the first to describe this relatively newly discovered condition (HIV-LA). It involves loss of facial soft tissues, leading to serious changes (other areas of the body such as buttocks and feet can be also

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affected) associated with generalised lipoatrophy, apparently triggered by the antiretroviral therapy. Many patients perceive it as a highly stigmatising manifestation of their HIV infection since loss of facial fatty tissue leads to a gaunt appearance and may lead to issues in work, social, and personal relationships. Another serious issue is that it has been reported that patients reduce their adherence to their antiretroviral medication regime in order to avoid this highly visible wasting effect, thus jeopardising their treatment (Collins et al., 2000). Whether by choice or as a result of trauma, it is clear that the challenges faced by any repair and regeneration (RR) process are many and complex. Although it is, of course, impossible to identify any one anatomical region as being the most demanding in this respect, the craniofacial region must surely be a strong contender. The long healing and repair process starts with the triage team before finally reaching the reconstruction surgical team. Even before the RR process can properly begin, there is a long and difficult road for the trauma patient. For non-medical specialists, even a quick perusal of books such as ‘Head, Face and Neck Trauma’ edited by Stewart (2005) or ‘Evaluation and Treatment of Orbital Fractures: A Multidisciplinary Approach’, edited by Holck and Ng (2006) brings home the enormity of the task, not to mention the costs involved. Bone and soft tissue injuries are not always of immediate priority when other potentially vision- or life-threatening injuries are involved. However, once the immediate medical aspects of the trauma have been satisfied then the multi-disciplinary team of reconstructive surgeons become involved. A multidisciplinary approach and multiple skills are required in the ultimate effort to restore both function and appearance requiring access to a range of accredited, well-defined and tested repair materials to be used in the reconstruction process. In order to develop clinically usable products, the reconstructive surgeons and indirectly the materials scientists and engineers, must become conversant not only with their own fields of expertise as well as with the anatomy and specific functions of the repair sites involved, but also with the healing process. In addition to fractures of the facial and orbital bones, injuries to eyes, nose, mandibular, and lachrymal systems and soft tissues such as cartilage, muscles, tendons, ligaments, and nerves may all have to be considered. A comprehensive coverage of all of these is beyond the scope of this chapter. Hence we shall focus on one aspect of facial reconstruction that has been less well addressed in the literature; namely an overview of some of the materials used to repair and regenerate soft tissue both in terms of fillers and in terms of materials used at the hard–soft tissue interface. We will focus on the current range of commercially available repair materials, some relevant current research aimed at improving these materials, and finally the next generation of repair materials and new RR strategies.

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Facial reconstruction

There is a plethora of useful books describing the relevant tissues as well as the complexity of their interrelationships. As a starting point an almost historic, compact and very readable text is ‘Anatomy of the Head, Neck, Face and Jaws’ by Fried. It provides the anatomical language necessary to form a deep understanding of all the tissue structures from bone and cartilage to nerves and muscles (Fried, 1980). From a clinical perspective the craniofacial skeleton is divided into upper and lower regions. The main function of the lower facial skeleton is occlusal and includes the masticatory structures of the mandible and maxilla, whereas the upper facial skeleton serves as a protective device for housing vital organs. In their chapter on ‘Strategies for Bone Substitution in Craniofacial Surgery’, Fialkov et al. (2000), in addition to describing the interacting forces involved between the muscles and facial bones, also specify the importance of the overlying soft tissues. The morphology of the face depends not only on the underlying bony structure but also on the subcutaneous fat or adipose layer that interfaces between the bones and skin. Hence, in any major trauma or malformation, the repair and regeneration of more than one tissue will be required. Soft tissue repair requirements are to some extent dependent on bone repair. Some of the important and relevant conclusions these authors arrive at are: in order to withstand the overlying soft tissue compression, bone substitutes, resorbable scaffolds and osteogenic carriers all require a degree of rigidity in order to be effective in the reconstruction and augmentation of the upper facial skeleton and secondly, the requirements of bone substitutes in the upper facial skeleton are different from other parts of the body (Fialkov, 2000). As will be discussed in the following sections, the same can be said for the requirements of many soft tissue substitutes that form the interface with the hard tissue. One exception is in cases where the soft tissue is repaired using subcutaneous injections of fillers such as autologous adipose tissue. One significant problem encountered in facial repair is the fact that autologous bone resorption, as well as the short to medium term lifetimes of many soft tissue replacement materials, lead to volume changes that are highly visible in the face. Hence, where the repair and regeneration of both soft and hard facial tissues is concerned, the search for the ‘ideal’ implant or material that will maintain both volume and contour has lead to the widespread use of alloplastic (see Section 3.3.2) materials. From the available literature it is clear that there are still aspects of many craniofacial repair procedures and the materials used that would benefit from further in-depth fundamental research. Before embarking on soft tissue repair, a brief description of the tissues involved is pertinent.

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3.2.1 Tissues at the bone interface Bone is a vascularised tissue consisting of cells and a mineralised extracellular matrix scaffold. The principal mineral component of the scaffold is carbonated hydroxyapatite which, although closely related to hydroxyapatite [Ca10(PO4)6(OH)2], is unique. The most abundant of the collagens present is collagen Type I, which is also the second highest component. It provides flexibility and structure to the tissue. Bone also contains collagen XI, V and III, as well as proteoglycans such as chondroitin, versican and syndecan, serum proteins, phosphoproteins and gamma carboxy glutamate (Siebel et al., 2006). The craniofacial skeleton is made up of two distinct bone types: membranous bone and endochondral bone. In membranous bone formation, which is responsible for the majority of the bones, ossification through direct mineral deposition into the extracellular produced matrix is followed by transformation of the mesenchymal cells into osteoblasts. Ultimately, this leads to the formation of the frontal, parietal, nasal, zygoma, maxilla and mandibular bones (Fig. 3.1). Endochondral bone formation, on the other hand, is responsible for the occipital bone, nasal septum and cranial base. Here the initial cartilaginous template is mineralised. Osteoclasts invade and are replaced by osteoblasts which ultimately lead to bone formation. As a result of extensive remodelling, at maturity there is minimal original skeletal bone and this has important implications for the development of the materials used in RR strategies. Cartilage is an avascular tissue and one of four mineralised tissues found in the body. It consists of an extracellular matrix comprised mainly of a three-dimensional hydrated network of collagen fibres in which are embedded chondrocytes and proteoglycans. Depending on the type of cartilage, it may or may not be mineralised. For example, articular cartilage contains about 60% collagen Type II and 40% Types I, VI, IX, X and XI (Walsh, 2006). The extracellular matrix (ECM) contains 4–7% (wet weight) proteoglycans (PGs), with aggrecan the most abundant. From a material and chemical point of view, cartilage is a hydrogel due to its high water content and hence it is capable of resisting pressure. The PGs interact with water which swells them, thus stabilising the tissue and imparting compressive stiffness. This contributes to the viscoelastic properties of cartilage. Because it is anaerobic it also has low oxygen consumption. Bone and cartilage are very different materials and the interface between them is an indistinct ‘calcified zone’. Although collagen Type II is the most abundant cartilage collagen, calcifying cartilage is enriched with collagen Type X and contains extracellular matrix vesicles. Currently, there are as many as 28 different collagen types described in the literature. During the bone mineralisation process, which is regulated by growth factors, cytokines and hormones, apatite deposition occurs onto the collagen Type

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I matrix (this being the most abundant bone collagen). According to one source, the distinction between bone and cartilage was already recognised by Aristotle, who separated fish into either cartilaginous or bony categories (Hall, 2005). However, it was not until the 1690s and the invention of the microscope that researchers such as Leewenhoek and Havers were able to investigate the intricate microstructures of bone and cartilage (Hall, 2005). Polarised light microscopy and scanning electron microscopy (SEM) studies made it possible to establish the architecture (de Visser et al., 2008). Adult articular cartilage has a zonal architecture which is determined by the alignment of its collagen fibres. The three zones of alignment are: the superficial zone (closest to the articular surface) where the fibres are aligned parallel to the articular surface, the radial zone (closest to the bone) where they are normal to the surface, and the transitional zone (between the superficial and radial zones) where a continuous variation in average fibre orientation is observed. In addition, the chondrocytes change morphology from flattened and more aligned in the calcified zone at the bone interface to rounder in the superficial zone. Muscles are organs of motion: folklore has it that although it takes 17 muscles to smile it takes 43 to frown. They are mainly composed of muscle cells, which in turn contain myofibrils. The hierarchical structure means that these contain sarcomeres which are composed of the proteins actin and myosin. In most muscles, all the fibres are oriented in the same direction, running in a line from the origin to the insertion. They usually connect two or more anatomical structures, one of which is capable of moving: bone and skin, two bones, two parts of skin, two organs, etc. The function of muscles is intrinsically linked to ligaments, nerves, tendons, aponeurosis and fascia: for example, these control the energy for the muscle contraction and its direction. All the muscles of facial expression are attached to skin; at least at the point of insertion. In addition, they are all superficial. Fried (1980) points out that they have the particular characteristic of coming in a wide range of sizes, shapes and strengths. As well as conveying emotions and displaying facial expressions, they fulfil many important functions such as closing the eyes and moving lips and mouth during mastication. Interestingly, the facial expression muscles work synergistically rather than independently. Hence, any repair is complicated by the fact that they are interdependent in many of their functions. Any material used to replace or repair bone, cartilage or skin will clearly have some relationship with the muscles connected to them, not to mention interfaces (whether lesser or greater) with the ligaments. Adipose tissue (AT) is a specialised, loose, non-fibrous connective tissue composed primarily of fat cells called adipocytes. Its most important role is in energy homeostasis. In mammals, fats, usually in the form of triglycerides, are stored either in white adipose tissue (WAT) or (to a lesser extent in

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adults) brown adipose tissue (BAT). WAT contains a single lipid droplet, while BAT contains numerous smaller droplets and contains many more mitochondria. The latter are responsible for the brown colour. BAT contains more capillaries due to its greater need for oxygen and is the main fat present in newborns. It is also responsible for generating body heat. Later, this is mostly replaced with WAT, although some BAT remains in the neck and intrascapula regions. WAT acts as an insulator and a protective layer around vital organs (think of the kidneys). In the face it fulfils another vital function where it forms a cushioning layer with interfaces between bone and skin helping to shape and contour the facial features. The skin which forms the external covering of the body is its largest organ, constituting 15–20% of its total mass and a surface area of between 1.5 and 2 m2 for the average human. The term ‘integumentary system’ is used when derivatives such as mammary, sweat and sebaceous glands are included. Skin is broadly categorised as thick or thin; a reflection on its location since it can vary from 1 mm to over 5 mm such as on the palms of the hands or soles of the feet. However, although these terms, which refer only to the epidermis, are really only of interest from a histological point of view, they are mentioned to illustrate the fact that skin can differ anatomically. As well as the two main layers – the epidermis and the dermis – there is a third underlying hypodermis fatty (subcutaneous fat) layer (see Fig. 3.3). This layer is highly significant in facial repair and reconstruction (RR) and is relevant to this chapter. The epidermis is composed of a keratinised, stratified squamous epithelium derived from the ectoderm. It grows continuously but, through the process of desquamation, maintains its regular thickness. One of its primary functions is as a protective barrier from the environment. The dermis is a dense connective tissue whose functions include imparting mechanical support and strength as well as thickness to the skin. Two distinct layers are clearly identified using light microscopy: the papillary and the reticular. Just beneath the reticular layer are found the layers of the hypodermis: lobules which contain varying amounts of adipose tissue separated by connective tissue septa, smooth muscle and, in some sites, striated muscle. This is the layer that insulates inhabitants of cold climates. But it is also the layer that varies greatly from individual to individual, bestowing on each of us a large part of our ‘identity’ based on how thin or not ‘quite so thin’ we are. This is particularly true where the face is concerned.

3.2.2 Organs of special senses: eye, nose and ear Vision, hearing, equilibrium and smell are all intrinsically linked to the anatomy of the head. While recognising the importance of all of these there

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Sebaceous gland

Skin

Subcutaneous fat (SF) Soft tissue Veins and arteries

(SF)

Muscle Soft tissue– bone interface

Bone

Bone

3.3 Tissue cross-section from skin to bone, including soft adipose tissue.

does seem to be some consensus that vision is the most important of all the senses and its protection is vital. Although the ‘eye’ itself can be considered as consisting of essential components such as the bulb and the accessory organs (muscles, eyelids and eyebrows, lachrymal system, etc.), where facial trauma is concerned the bony orbit is critically important because this is where the eye is housed and protected. Broadly speaking, it consists of the orbital rim which is relatively thick and the orbital walls which are much more fragile. A recent book ‘Biomaterials and Regenerative Medicine in Ophthalmology’ edited by Prof. Traian V. Chirila, Queensland Eye Institute, Australia (Chirila, 2010) covers many aspects of the repair of this particular system. The external nose is mainly composed of bone, cartilage, skin and mucosa. Nasal fractures are the most common facial trauma (∼40%) and for whole body trauma they rank third highest overall after wrist and clavicle (Dev, 2008). The history of medicine, and more particularly the materials used to repair the body, make fascinating reading, even if their historical accuracy is sometimes somewhat doubtful. One of the most interesting is the story

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of Tycho Brahe’s artificial nose (1566) (Van Helden, 1995). After losing his nose (or part of it depending on which account one reads) in a duel, he had a gold/silver prosthesis fitted using an ‘adhesive balm’ to keep it in place. This incident apparently kindled his interest in medicine, although he is best remembered for his astronomical discoveries. Another well-documented case of nasal repair is found in the Edwin Smith Papyrus, which is the only surviving copy of an ancient Egyptian book on trauma surgery. It consists exclusively of cases beginning with the head and working down. According to one report, of the 48 surviving case-reports only one resorted to ‘magic’ (Wilkins, 1964). Another case described involved the ‘repositioning of the deviated nasal bones’. Treatment included ‘the use of internal splints, firm external splints and dressings made from linen and grease and honey’. Further fascinating reading can be found in Lascaratos et al. (2003) ‘From the Roots of Rhinology: The Reconstruction of Nasal Injuries by Hippocrates’. Returning to the present, Oeltjen and Hollier (2005) give an excellent overview of ‘Nasal and Naso–Orbital Ethmoid Fractures’ in ‘Head, Face and Neck Trauma’ edited by Stewart (2005). In addition to hearing, the ear fulfils two other functions: equilibrium and cosmesis. The external ear consists of two major parts; the auricula or pinna and the external auditory meatus (Fried, 1980). Even focusing on the pinna and cosmesis issues, its repair is by no means trivial. It is composed of soft tissues, with cartilage being dominant. The chapter by Chang on ‘Auricular Trauma’ gives a good overview of typical trauma, treatment options and strategies (Stewart, 2005). Typically, microsurgical techniques have been used since the 1980s to reattach partially avulsed pinna. In cases where the total external ear is missing from birth, surgery (four operations over 2 years), is possible and most successful when the patient is 6–8 years old. When the loss is due to radical cancer surgery, amputation, burns and/or congenital defects, then auricular prostheses are available (UT Southwestern, 2009).

3.3

Materials used in traditional interfacial repair

Since in this chapter we are focusing principally on repair of the tissues underlying the skin and interfaced between it and the facial bones, as shown in Fig. 3.3, we now need to consider the materials and current approaches to facial soft tissue repair. There is a host of products used for soft tissue augmentation or replacement. Table 3.1 lists a selection of currently available materials. Our focus is on the three main classes of materials used either as fillers or for other soft tissue replacement applications. Depending on the genre of journal (or research), the descriptive words used in the classification of materials can vary. For example, the terms temporary, semipermanent or permanent are often used, but in biomaterials science the

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5% PAA injectable gel 2.5% PAA injectable gel

Beads suspended in bovine collagen gel Silicone Silicone Silicone

PMMA Silicone

PTFE PTFE-graphite PTFE-aluminium oxide Expanded PTFE Expanded PTFE tubing Expanded PTFE tubing Expanded PTFE with silver and chlorhexidine (antibacterial) Expanded PTFE dual porosity Expanded PTFE saline filled

Description/use

Polyacrylamide gels (PAAG)

Biostables PTFE

Material

ArteFill Silastic Silikon 1000 Adatosil 5000

Aquamid Eutrophill

Teflon Proplast I Proplast II Gore-Tex SoftForm UltraSoft MycroMesh Plus Advanta Fulfil

Trade name

Table 3.1 A selection of polymeric materials used as soft tissue fillers or in soft tissue repair

Artes Medical Dow Corning Alcon Labs Bausch & Lomb

Contura International Lab Procytech

Atrium Medical Products Evera Medical

Dow Chemical/Dupont Vitek Vitek WL Gore & Associates Tissue Technologies Tissue Technologies WL Gore & Associates

Manufacturer

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Rooster-derived hyaluronic acid Rooster-derived hyaluronic acid Bacterial or non-animal stabilised Bacterial or non-animal stabilised Bacterial or non-animal stabilised Bacterial or non-animal stabilised

Gelfilm (absorbable gelatin film)

Hyaluronic acid

Gelatin

hyaluronic hyaluronic hyaluronic hyaluronic

Bovine collagen Bovine collagen Cross-linked bovine collagen Human collagen Human collagen Cross-linked human collagen Human harvested autologous cells Human cadaver allogeneic collagen Human harvested autologous cells Porcine collagen

Stimulatory filler

Bioresorbables Polylactide

Naturally derived Collagen

Description/use

Material

Table 3.1 Continued

acid acid acid acid

Gelfilm

Hylaform Hylaform plus Restylane Perlane Prevelle Juvederm

Zyderm I Zyderm II Zyplast Cosmo Derm I Cosmo Derm II Cosmo Plast Isolagen Dermalogen Autologen Evolence

Sculptra/NewFill

Trade name

Pharmacia & Upjohn Ophthalmic Division

Allergan, Inc. Allergan, Inc. Medicis Medicis Mentor Allergan, Inc.

Allergan, Inc. Allergan, Inc. Allergan, Inc. Allergan, Inc. Allergan, Inc. Allergan, Inc. Fibrocell Science, Inc. Collagenesis, Inc. Collagenesis, Inc. ColBar LifeScience Ltd

Sanofi Aventis

Manufacturer

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terms ‘bioresorbable’ (degrades in vivo), ‘non-biodegradable’ (permanent) and ‘naturally derived’ (sourced from living organisms) are preferred. Recently, it has been suggested by one surgeon that another approach could be to consider them ‘according to the reactions they induce within human tissue’ (Nicolau, 2008). Seeking to control the rate at which bioresorbable materials degrade (or erode) in vivo is one that attracts an enormous amount of research since the processes involved greatly influence vascularisation and/or new tissue growth, both of which in turn influence the extent of tissue regeneration. It should be pointed out here that, in general, naturally derived repair materials (e.g. adipose tissue) degrade very fast and often in an unpredictable manner. In contrast, the degradation rate of synthetic bioresorbable materials can often be tailored for a specific application. Klein and Elson (2000) listed the requirements of a soft-tissue augmentation material as being: • • • • • • • • •

potentially of high use cosmetically pleasing have minimum undesirable reactions have low ‘abuse’ potential or abuse not leading to significant morbidity rates non-teratogenic non-carcinogenic non-migratory capable of providing predictable, persistent, reproducible correction FDA [presumably also other regulatory bodies] approved if non-autologous

Because of the dominance of the ‘facial cosmetic surgery’ market, even a superficial search reveals that there is a plethora of injectable materials (or fillers) described in a huge variety of journals which are used to ‘improve’ facial appearances or rejuvenate ageing features. Comprehensive lists can be found in various reviews such as ‘Soft Tissue Augmentation: A review’ (Fernandez and Mackley, 2006) and ‘Long-lasting and Permanent Fillers: Biomaterial Influence over Host Tissue Response’ (Nicolau, 2008). This ‘market’ and the controversies surrounding its regulation in different countries are not of concern here. However, since many of the materials used also play an important role in the repair of facial defects, scars and trauma injuries, these are being included. As Klein and Elson (as well as many others) pointed out currently, there are no repair materials that fulfil all the desired criteria (Klein and Elson, 2000; Nicolau, 2007). So it is not surprising that it is also generally recognised that the ‘perfect’ filler has still to be found. In the words of Hirsch in her recent review on ‘Soft Tissue Augmentation’ (Hirsch and Cohen, 2006), ‘even a judicious selection of the

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perfect product for a given indication in a particular patient does not guarantee a perfect outcome’. Time and time again in the literature there are warnings about using particular products without having the necessary knowledge and expertise. However, one thing any surgeon or material scientist will agree on is that this is an ambitious ‘wish list’ for any material to possess! We shall now discuss the most frequently used materials in more detail. Table 3.1 lists a selection of soft tissue augmentation and filling materials. This unprejudiced list is by no means complete but is an example of representative materials for each of the classes to be discussed. We have included autologous materials such as fat and collagen but not botulinum toxin.

3.3.1 Naturally derived materials Although two of the most frequently used naturally derived materials, ‘autologous fat tissue’ and ‘collagen-based materials’, lie outside the main focus of this chapter, we include a short overview of each. The oldest and most frequently used autograft is in fact ‘autologous fat tissue’. Neuber first used it in whole graft form in 1893 (Neuber, 1893). Since the 1920s it has been used in injectable form, but with the advent of liposuction its use has greatly increased. There is much controversy surrounding its efficacy, in particular its longevity, and this is reflected even in the titles of some recent reviews: ‘Autologous Fat Transfer for Facial Recontouring: Is there Science behind the Art?’ (Kaufman et al., 2007) and ‘Fat Grafting: Fact or Fiction?’ (Calabria and Hills, 2005). It is used as a non-permanent material to fill small defects and for scar repair. Because of the large degree of resorption (30–60%), substantial over-correction is necessary. The harvesting and tissue preparation techniques used are well documented. Despite the many disadvantages, its use has persisted because its autologous nature means biocompatibility is not an issue. Naturally derived materials come from a variety of sources including autologous (human) and allogenic sources such as bovine, porcine, avian, as well as bacterial. Collagen was one of the first naturally derived filler materials given FDA approval (Sarnoff et al., 2008). Bovine-derived collagen is one of the most popular injectable soft tissue fillers and is generally considered the gold standard against which all other materials are measured (Homicz and Watson, 2004; Klein and Elson, 2000). This is not really surprising, since collagen itself is a main structural element of both hard and soft tissues in mammals. Bovine collagen-based materials have been used for over 20 years and the main disadvantages associated with their use are well documented: a propensity for hypersensitivity reactions and limited efficacy lifetimes. The need for skin testing requires multiple visits and hence delayed treatment times. Abscess formation, tissue necrosis and

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granulations as a result of foreign body reactions have all been reported (Hanke et al., 1991). Some of the many autologous and allogenic collagenbased materials that have been developed over the years are expensive, customised products. One such commercially autologous product is Isolagen® which contains the patient’s own living dermal fibroblast cells (Isolagen, 2007). More recently, in their review, Homicz and Watson attest that hyaluronic acid-based materials such as Restylane® and the Hylaform® range of products appear to be a ‘safe alternative’ to bovine collagen-based materials (Homicz and Watson, 2004). However, even though hyaluronic acid structures are believed to be identical across species (for example it is isolated from rooster combs and streptococcus), traces of hyaluronic acid-associated proteins can cause allergic reactions (reactions currently reported as <1%) (Lowe et al., 2005). The rationale for the use of naturally derived polymers, e.g. collagen (Table 3.1), particularly when they are present in the patient’s own system, appears to be logical. These polymers usually degrade in vivo enzymatically but many are also susceptible to hydrolysis. The degradation by-products are usually disposed of, or recycled, by the body through normal metabolic pathways. Furthermore, because of the chemical similarity between these polymers and extracellular matrix components already present in tissues, biocompatibility and integration would be expected to be enhanced. However, these polymers have very fast and uncontrollable degradation rates in vivo. In addition, in order to produce enough material, the crude polymer is usually sourced from a different species to the patient. As a result, there is concern over, not only disease transmission, but also the variable quality of these polymers, which often differs between batches.

3.3.2 Bioresorbable and non-biodegradable materials According to Eppley (2003) and others, the concept of an ‘alloplastic’ material is synonymous with the term ‘synthetic’ and means it originates from a non-biological source. When used in the medical literature, the term alloplastic covers manufactured materials from ‘non-organic, non-human, nonanimal sources’, and encompasses ceramics and metals as well as plastics (which to materials chemists and materials engineers means polymers). We are concentrating our discussion in this chapter on polymeric materials. Broadly speaking, polymeric biomaterials are divided into either nonbiodegradable (such as polytetrafluoroethylene (PTFE)) and bioresorbable polymers (such as the poly(lactic acid)-based Lactosorb®). Because different disciplines such as tissue engineering and materials chemistry use different definitions of the terms biodegradable, bioresorbable, bioabsorbable and bioerodible, some confusion has arisen in the literature. So although

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the term biodegradable is the one most frequently used in its broadest sense, it is pertinent to define these terms because they can be important when discussing the chemical and physical properties of the kinds of polymers discussed in this review. Vert’s definitions, given below for solid polymeric materials, are generally accepted by the materials and tissue engineering communities (Albertsson and Varma, 2003; Vert et al., 1992). •

Biodegradables break down due to macromolecular degradation. There is in vivo dispersion of the fragments/by-products but no proof of elimination from the body. • Bioresorbables show bulk degradation and further resorb in vivo: i.e. the original foreign material and its breakdown products can be shown to be eliminated through the body’s natural pathways. • Bioerodibles show surface degradation and further resorb in vivo. Total elimination of low molecular weight by-products is inherent. • Bioabsorbables can dissolve in body fluids in the absence of polymer chain cleavage or molecular mass loss, such as in the slow dissolution of water-soluble materials. In addition to the many well-documented tissue RR applications for the materials shown in Table 3.1 another potential application (already mentioned in Section 3.1) which is very relevant to our topic is in HIV-LA surgery. However, some alloplastics are more used than others and although many are approved for use in Europe, they are not always FDA-approved. Hence we shall limit our discussion to two examples which are well studied and documented: PLA (bioresorbable) and ePTFE (non-biodegradable). PLA belongs to a class of compounds known as the poly(α-esters). Due to their good biocompatibility and ability to be bioresorbed, aliphatic poly(α-esters), such as polylactide (PLA), polyglycolide (PGA), and their copolymers have been studied for biomedical purposes since the 1960s (Albertsson and Varma, 2003). The first commercially available product launched in 1962 was a polyglycolide suture called DexonTM (Tyco Healthcare Group, CT, USA). Since then, PLA has manifested itself in many forms for use in medical devices and in facial surgery applications. According to Moyle, PLA is the only one which has specific FDA approval for treatment of HIV-LA (Moyle et al., 2004, 2006). In a series of papers, he and others (Valantin et al., 2003; Burgess and Quiroga, 2005; Lafaurie et al., 2005; Barton et al., 2006; Cattelan et al., 2006; Lam et al., 2006) describe and evaluate injectable forms of PLA for use as a bioresorbable filler with a lifetime of 1–2 and sometimes up to 3 years. New-Fill® has been used since 1999 in Europe (Fernandez and Mackley, 2006). According to Hirsch, Sculptra®, as it is known in the USA, was fast-tracked in 2004 for approval by the FDA for use in facial lipoatrophy (Hirsch and Cohen, 2006).

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One of the problems associated with the use of injectable PLA suspensions is the formation of subcutaneous papules and granulomas. As a result of various studies, (Valantin et al., 2003, Woerle et al., 2004) suggestions as to how to avoid these side-effects have been made. They are chiefly concerned with practical issues such as depth of injection, frequency of reapplications, and post-injection strategies (ice packs). However, from a fundamental biomaterials science perspective, the fact that these fillers are administered as suspensions of PLA microparticles means that overall, they have a large surface area and the body’s immune response will be to reject or encapsulate each and every one of the ‘foreign body particules’ injected. Similar side effects have been observed for non-resorbable fillers administered as microparticular suspensions, e.g. polymethyl methacrylate (PMMA) (Lemperle et al., 2003; Homicz and Watson, 2004). In his review, Nicolau (2008) addresses many relevant issues including the issue of particle size and surface area. This and other references cited therein merit reading as there appears to be some difference of opinion on what constitutes the ‘ideal’ particle size. For example, in their 2002 paper, Morhenn et al. describe how PMMA particles less than 20 μm ‘appear’ to have been phagocytosed by U-937 cells (human macrophage-type cells), while particles greater than 40 μm were not. This is in contrast to collagen-coated PLA microspheres 0.6–60 μm, which do ‘appear’ to be all phagocytosed. Results for PMMA microspheres with Langerhans cells differ to some previously published results (Reis e Sousa et al., 1993) and it was concluded that ‘the different chemical nature of the microspheres and the type of Langerhans cells cultures may explain the differing results’. On the other hand, it is well accepted that smooth, regularly shaped particles produce a much smaller inflammatory response than irregularly shaped ones (Nicolau, 2008 and refs therein). The many different commercial filler products all have subtle composition/ structural differences and hence it becomes virtually impossible to predict the extent of an inflammatory response. Another issue with ‘microsphere suspension fillers’ is the fact that sometimes the microspheres can migrate and this, of course, can lead to a new set of inflammatory responses. Again, it appears to be up to the individual surgeon to decide whether or not to use a particular filler for a specific application. For example, Nicolau (2008) states that he does not use the PLA product NewFill® in the lips because they are known to displace and form unsightly nodules.

3.3.3 The polytetrafluoroethylenes Polytetrafluoroethylene (PTFE) is a fully fluorinated unbranched polymer with a carbon backbone, analogous to the hydrocarbon polymer polyethyl-

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Biointegration of medical implant materials F F F

F

F

F

F

F

F

F

F F

F

F

F

F

F

F

F

F

F

F

F F

F F F

F F

F C

F

F

F

F C

F

F

3.4 The helical structure of polytetrafluoroethylene.

ene (PE) (Sperati and Starkweather, 1961; Scheirs, 1997; Bunn and Howells, 1954). However, the conformation of PE is the common planar zig-zag conformation whereas PTFE is helical, Fig. 3.4 (Sperati and Starkweather, 1961; Scheirs, 1997). The comparatively large size of the fluorine atoms and the mutual repulsion between the adjacent fluorine atoms causes the polymer chains to twist to form compact helical rods with 13 fluorine atoms per 180° helical turn (Bunn and Howells, 1954; Drobny, 2000). Because it is easier to deform the C–C bond angle than stretch the C–C bond, each main-chain bond is rotated 20° from the next. This conformation creates a stiff rod shape, with the inner core of carbon completely encased within a sheath of electronegative fluorine atoms. This sheathing makes the rod conformation extremely stiff by inhibiting bending of the carbon backbone (Schiers, 1997). Consequently, the polymer chains are, of necessity, parallel packed and this, combined with the high electronegativity of the outer sheath of fluorine atoms, allows the rods to slide past one another, culminating in the tendency of PTFE to cold flow (creep) (Schiers, 1997). It is likely that this phenomenon gives rise to the issues (discussed more fully later) of concern in implant surgery; namely shrinkage, migration, and micromovement. The C–F bond is the strongest chemical bond found in polymers and confers on PTFE an exceptional chemical resistance (Sperati and Starkweather, 1961; Dargaville et al., 2003). It exhibits good electrical insulating abilities, a property that contributes to many commercially important

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applications. It is tough, non-adhesive and has anti-frictional properties combined with extreme hydrophobicity (Sperati, 1986; Gore, 1976a). The unique chemistry of PTFE gives it its unusual qualities of being both thermally and chemically inert. Although it is its inert nature that first made PTFE attractive as a biomedical material, it is also this property that leads to its lack of surface adhesion in vivo. This may necessitate aggressive suturing or anchoring in some applications, particularly in some facial ‘tight pockets’. PTFE’s unique properties and versatile product range means that it has been used in medicine since the mid-twentieth century. The first use of the non-expanded material was as an artificial heart valve, and later it became popular as a vascular graft replacement (Kannan et al., 2005). Voorhees and co-workers highlighted the importance of porosity in arterial grafts (Voorhees et al., 1952). Until then, PTFE had been woven into a textile graft form. This was not ideal for the highly crystalline PTFE, which meant sacrificing some of its excellent features and led to unravelling in the hostile in vivo environment. Expanded PTFE (ePTFE) is initially fibrillar and then expanded, and has proven more favourable for medical applications than the non-expanded versions. This is not only because it is non-biodegradable but also because it has a low propensity to trigger thrombosis. These expanded materials have been described as having a ‘microstructure characterised by nodes interconnected by fibrils’ (Gore, 1976a,b), as shown in Fig. 3.5. The size of the nodes and the character of the fibrillar structures are dependent on the conditions used in the expansion process. Currently, ePTFE has a porosity of 70 to 85 percent, with pore sizes ranging from 0.5

3.5 SEM image of ePTFE showing nodes and interconnecting fibrils.

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to 100 micron, depending on the source of the material. The porous microstructure allows ingrowth of tissue, which is essential for implant fixation, acceptance, and positive outcomes. Expanded PTFE tubing produced as Gore-Tex was first used clinically in portal vein, thoracic vena cava, inferior vena cava, and external iliac veins in 1972 (Soyer et al., 1972). Since then there has been a plethora of reports on the use of ePTFE, particularly Gore-Tex. Kannan et al. (2005) provide an excellent review of both models and clinical outcomes of prosthetic bypass and microvascular grafts reported in the literature from 1984 to the present. Reports indicate that both ePTFE and PTFE perform on a par with other synthetic grafts. However, they recommend their restriction to high flow vessels. Smaller diameter grafts tend to fail due to occlusion caused by thrombosis and intimal hyperplasia (Bezuidenhout and Zilla, 2004). Guidoin et al. (1993) performed chemical analyses on 79 explanted arterial prostheses specimens removed after 6.5 years from human patients and concluded that ePTFE is stable for at least this duration in vivo. PTFE’s low friction coefficient has meant that it is an excellent coating material for devices that require ease of placement or ease of movement; for example, catheters and artificial joint coatings. PTFE-coated catheters were introduced during the late 1960s as their low friction properties minimise patient discomfort (Lawrence and Turner, 2005). The main longevityrelated problem in total joint replacement is wear; particle-induced osteolysis around the acetabulum (cup) component. It has been demonstrated that using an ePTFE membrane as a physical seal prevents loosening and wear: subsequently osteoblasts were found within the ePTFE seal (Bhumbra et al., 2000). Periodontal applications were the first to exploit the principles of guided tissue regeneration (GTR) (Zhao et al., 2000), later described as guided bone regeneration (GBR) for applications involving bone defects (Dahlin et al., 1991). Many periodontal concerns and treatments, such as periapical bone resorption (occurring at the base of the tooth apex), are applicable to GBR treatments elsewhere in the body, including maxillo- and cranio-facial sites. In such cases, bone healing often fails to occur or occurs improperly. Large voids can occur either congenitally or through trauma, cysts, or natural resorption of the bone. Regeneration of bone at non-union sites is dependent on patient age, bone structure, vascularisation, the soft-tissue environment as well as the size of the defect (Dahlin et al., 1994). Regeneration is often impaired by rapid ingrowth into the wound site by connective tissue from the surrounding areas. Extensive research has taken place in the area of GBR using ePTFE as a non-resorbable membrane. In fact, in their review on the use of membranes for bone healing and neogenesis, Linde et al. (1993) state that virtually all investigations published on the promotion of bone regeneration by

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the membrane technique have utilised ePTFE membranes. There has been extensive experimental work using a variety of animal models from mice (Kidd et al., 2002) and rats (Dahlin et al., 1994; Kidd et al., 2002; Hagerty et al., 2000) to sheep (Paavolainen et al., 1993), pigs (Wiltfang et al., 1998) and monkeys (Hürzeler et al., 1998). These are only a small sample of both the model type and studies performed. One of the very early studies on rabbits is one in which Neel was the first (as far as we are aware) to suggest that these materials could be used in cranio-facial applications (Neel, 1983). Furthermore, clinical periodontal radiographic studies were performed by Lorenzoni et al. (1999) to determine the regenerated bone response over 2 years to functional loading in vivo for 82 patients fitted with ePTFE GBR membranes in conjunction with prosthetic teeth. The regenerated bone appeared able to withstand functional loading. Periotest values were done at 6 and 24 months and revealed stable peri-implant conditions and sustained osseointegration. Dahlin et al. (1998) tested the histological morphology of the ePTFE/tissue interface in oral implant patients. In their study, the membrane was penetrated by fibrous connective tissue which also separated the membrane from the bone. They concluded that fine micro-movements of the implants might have caused this negative response. It is interesting to find that the results, both quantitative and qualitative, across all these studies showed no significant differences between the species. Expanded PTFE is used as a soft tissue augmentation material for the face in the anterior nasal spine, glabella area, lips, malar area, mandible, mentum, mid-face region, nasal dorsum, nasolabial folds, and premaxilla (Mercandetti et al., 2008). It can be processed into a variety of forms, three-dimensional shapes, sheets, tubes, etc., which demonstrates its functional versatility. In particular, it was the ability to create these three-dimensional shapes that made ePTFE so valuable in facial softtissue suspension, augmentation, and correction (Panossian and Garner, 2004). Early PTFE included Proplast I. This was a black PTFE–graphite composite and was initially designed as a coating material. It was replaced by Proplast II, a white PTFE–aluminium oxide composition which could be easily cut and shaped. The use of both Proplasts in these early applications was discontinued due to the many reports of gross malfunction. In some cases these materials had been used in the temporomandibular joint (TMJ) and hence were put under considerable load. Their failure included significant wear, fragmentation, and perforation. The particulates formed in the immediate anatomical area triggered a foreign body response, thereby causing bone degeneration (including the glenoid fossa, mandibular condyle and erosion into the cranial space). In some cases, the particulates were found to migrate to the lymph nodes where they could cause permanent

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hearing damage, chronic pain, loss of masticatory function and motion (Food and Drug Admin, 1991). Until recently, the originators of PTFE marketed as Gore-Tex® (W L Gore and Associates, Flagstaff, Ariz.) were producing this material in a variety of products suited to facial soft tissue repair. These included solid tubes as well as patches in both preformed and customised shapes (Panossian and Garner, 2004; Chandler-Temple et al., 2008). Interestingly, although this product had been used since the early 1970s, it was not FDA approved for use for facial reconstruction and augmentation until 1993. In 1997, the FDA approved the use of the tubular implant SoftForm® (Collagen Corporation, Palo Alto, Calif.). SoftForm® was used to treat deep facial furrows, such as nasolabial folds, as well as lip augmentation. The tubular SoftForm® and the solid strand Gore-Tex® used for the same purpose have similar porosities of 5 to 30 μm (Truswell, 2002). SoftForm® was eventually replaced by tubular UltraSoft® (Tissue Technologies Inc., San Francisco, California) in 2001. Cox (2005) reported on the advantages and disadvantages of permanent implants (ePTFE) vs non-permanent (fillers). The permanent implants are easy to remove if they turn out to be mal-positioned or if subsequently desired by the patient. However, they have limited filling capability and can feel undesirably firm to the touch. They can form elevated regions, extrude, contract, and migrate. The issue of shrinkage has been observed experimentally (Chandler-Temple, 2007). To address some of these issues (migration, shrinkage, and softness), Advanta® (Atrium Medical Corporation, Hudson, NH, and Ocean Breeze Surgical, Amherst, NH) was developed and has been available since 2001 (Panossian and Garner, 2004; Redbord and Hanke, 2008). It consists of non-hollow strands of dual porosity: outer and inner core porosities of 40 μm and 100 μm respectively. In late 2006, after W L Gore withdrew their Gore-Tex SAMs series, Advanta® was the only product of this type available commercially (Redbord and Hanke, 2008). They continue to market and develop their MycroMesh Plus implant which incorporates silver and chlorhexidine into the ePTFE matrix, producing a material with similar qualities to Gore-Tex®. However, they report a greatly reduced infection rate (Malaisrie et al. 1998; Panossian and Garner, 2004). Throughout the literature, there is an emphasis on the need for tissue in-growth to assist in the biointegration and fixture of these implants. The nature of the tissue in-growth needs to be carefully considered. Ideally there should be enough specific cellular in-growth to anchor the material, without the formation of a fibrous capsule because this not only hinders the appropriate cellular integration, but also leads to micro-motion and ultimately more serious consequences. This brief review of the products available for facial soft-tissue augmentation demonstrates the ongoing development of new approaches to engi-

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neering materials in order to improve both material properties and implant outcomes. Tissue in-growth has been improved as a result of pore size engineering. Tactility as well as antibacterial properties have also been improved by the customised engineering of this exceptional material. However, these developments have not addressed a most critical aspect of the implant; namely, its surface. It is the material surface that first comes into contact with the patient’s internal biological milieu and the first response of the body’s immune system is to the surface. Hence, the surface chemistry of implants and not just their gross morphologies should be considered.

3.4

Surface modification of facial membranes for optimal biointegration

As described in the previous section, ePTFE membranes have been used as soft tissue replacement materials at the soft tissue–bone interface. However, due to the inert nature of ePTFE it does not integrate sufficiently with the underlying bone and thus a number of research groups have been investigating methods to improve the bonding strength at this interface. Because PTFE is both thermally and chemically inert, its surface needs to be exposed to high energy processes in order to change its properties. Some of the approaches that have been used include ion implantation (Colwell et al., 2003), ArF excimer laser irradiation (Sato and Murahara, 2004), plasma activation using various carrier gases such as H2O or N2 (Oehr et al., 1999), O2 (Kang et al., 2001), and Ar (Hsueh et al., 2003), electron beam pre-irradiation (He and Gu, 2003), and 60Co gamma pre-irradiation (Mazzei et al., 2002). In addition, simultaneous grafting can be performed using plasma polymerisation (Kühn et al., 2001), proton beam irradiation (Mazzei et al., 2003), as well as 60Co gamma irradiation (Dargaville et al., 2003). Control of the induced changes can be achieved through judicious choice of reaction conditions, e.g. for gamma irradiation-induced grafting these include monomer, dose rate and solvent. The surface properties of the modified polymers can differ substantially from those of the parent polymers, a fact that can be advantageously exploited to produce new materials with specific properties. Of the approaches listed above, the use of pretreatments (e.g. plasma or laser) prior to applying a calcium phosphate coat, and gamma irradiation-induced grafting have been explored for the improvement of ePTFE membranes used in facial applications. We will now describe these in some detail.

3.4.1 Calcium phosphate coatings It is well established that a coat of a calcium phosphate phase, preferably hydroxyapatite (HAP, Ca10(PO4)6(OH)2), on the surface of a biomaterial

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increases its ability to integrate with bone tissue in vivo. This has been extensively used for the fabrication of orthopaedic metallic implants and has been shown to greatly improve the stability of the interface (Ratner et al., 2004). The method used to coat metallic implants, however, is not easily transferable to polymeric substrates. This is due to the fact that high reaction temperatures above the polymer melting point are required. A second issue when adding a coat to a substrate is the need for a strong interface between the polymer implant and the coating material itself. Since PTFE is highly hydrophobic and has a very low friction coefficient, and HAP is highly hydrophilic, the resulting interfacial bond between them is very weak. In order therefore to produce a material with a well-bonded mineral coat on ePTFE membranes, it is necessary to first pre-treat the membrane to decrease its hydrophobicity. Early work in this area was conducted by Hontsu et al. (1998), who used an ArF excimer laser deposition technique to produce thin HAP films on PTFE substrates that had been chemically pre-treated with a sodium– naphthalene complex. The chemical pre-treatment introduced a mixture of oxygen-containing functional groups onto the polymer surface (e.g. –OH, –COOH), thus decreasing the hydrophobicity of the polymer. The HAP coat deposited at 310 °C was highly amorphous but increased in crystallinity upon annealing. The tensile strength of the interface was found to be 6 MPa. This was an improvement by an order of magnitude compared to a coating applied to a non pre-treated PTFE substrate. Jansen and his research group used radio frequency magnetron sputtering to deposit HAP onto PTFE substrates (Feddes et al., 2004a,b,c). This is a technique for depositing a thin HAP coating that can also be applied to polymers since relatively low temperatures can be used. The coating composition was found to be dependent on a number of parameters such as discharge power (low power needed for polymers), sputter gas composition, gas pressure, and position of the sample in the chamber. In addition, the nature of the substrate and the thickness of the coat were also found to affect the coating composition. In their study on virgin PTFE (Feddes et al., 2004a) they found that the amount of calcium in the coating could be controlled by the Ar gas pressure: low levels were introduced at high pressures. In addition, UV irradiation caused detachment of F atoms from the polymer substrate. This was found to lead to a high Ca/P ratio (up to 5) as a result of calcium combining with the escaping F atoms. The result was the formation of CaF2 in the coating. At the same time, it was proposed that the removal of P occurred due to the formation of volatile PF3 molecules. Despite this, the interfacial bond between the coating and the PTFE substrate was very strong (5.8 MPa) and displayed cohesive failure (i.e. within the PTFE substrate) during adhesion testing (Feddes et al., 2004b). In a subsequent study, various pre-treatments including O2 plasma pre-

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treatment, Ar+ ion gun treatment, and deposition of a Ti interlayer prior to coating deposition were evaluated (Feddes et al., 2004c). It was found that the presence of a Ti interlayer dramatically improved the adhesion of the coating (evaluated using the diamond scratch test). The other pre-treatment methods, however, did not improve the adhesion of the coating layer onto the PTFE substrate. The reason proposed for the observed optimal adhesion for the Ti pre-treatment was the ability of Ti to protect PTFE from UV irradiation-induced degradation. In addition, it was proposed that chemical bonding between the Ti interlayer and the PTFE substrate was an integral part of the system produced. Recently, Yamaguchi et al. (2008) used an electrophoretic deposition technique. They applied an insulating photo-resist mask to deposit patterns of wollastonite (CaSiO3) onto a porous ePTFE substrate. Subsequent biomimetic HAP growth was achieved by immersion in simulated body fluid (SBF), leading to mineral nucleation at the wollastonite sites. The resulting HAP patterns were 200 μm in size and displayed distinct boundaries. To date there has been no study reported on the adhesive strength of these coatings. Although all the techniques described here are promising as methods for improving the interfacial bond of ePTFE with the underlying bone, much work still needs to be done in order to assess the biological response to the introduced coatings.

3.4.2 Gamma irradiation-induced grafting Grafting techniques have been receiving increased attention in the area of biomaterials science due to their versatility. Simultaneous radiation-induced grafting involves the immersion of a polymer substrate in a monomer or monomer solution with simultaneous irradiation of the grafting system. This method produces radicals in the solution in the first instance and subsequently in the polymeric substrate (Wentrup-Byrne et al., 2005). The advantages of this method include its simplicity of execution, the versatility in tailoring the grafting outcome, and the inherent formation of homopolymer which protects the substrate from excessive radiation degradation. Conversely, the formation of excess homopolymer may hinder efficient graft copolymerisation and can thus be a disadvantage. Low doses are usually chosen, due to the tendency of PTFE to undergo radiation-induced chain scission (rather than crosslinking). Chapiro’s seminal works studying radiation grafting of styrene and methyl methacrylate onto PTFE showed the remarkable propensity of the monomers to swell the substrate and hence lead to homogeneous grafting throughout (Chapiro, 1962). The mechanism for this phenomenon became known as the ‘grafting front mechanism’ and has since been widely studied.

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It occurs when initial grafting on the substrate causes changes at the surface through swelling that subsequently allow the grafting to proceed gradually into the bulk polymer (Hegazy et al., 1992). Since these early advances, many studies have been performed grafting various monomers onto PTFE (Dargaville et al., 2003). There are, however, only a few reports of radiation graft polymerisation onto ePTFE using monomers such as polyethylene oxide (Park et al., 2000), acrylic acid (Grøndahl et al., 2003), and phosphonates (Chandler-Temple et al., 2010a; Suzuki et al., 2005; Wentrup-Byrne et al., 2005; Grøndahl et al., 2002). Whereas the first of these studies aims at improving the non-fouling properties for use of ePTFE in blood vessels, the remaining works investigate the surface improvement of ePTFE membranes for facial applications. The choice of acrylic acid, and in particular phosphate-containing monomers, for grafting onto ePTFE facial membranes originated from earlier studies that showed that these functional groups improve both the HAP nucleation process and growth, and therefore bone-bonding ability was inferred (often referred to as bioactivity). Indeed, the growth rate of apatite formation has been shown to decrease with the functional group in the order PO4H2 > COOH >> CONH2 ⬵ OH >> NH2 >> CH3 ⬵ 0 (Tanahashi and Matsuda, 1997). In addition, grafting phosphate-containing monomers onto polyethylene biomaterials resulted in doubling the amount of HAP growth on the modified material in SBF whereas a smaller effect was observed for the acrylic acid grafted material (Tretinnikov and Ikada, 1997). Subsequent in vivo studies on the phosphate-modified polyethylene biomaterial showed a significantly enhanced interface of the implant surface with the newly formed bone. This was attributed to the phosphate groups providing nucleation sites for HAP growth (Kamei et al., 1997). The initial studies on modifying the surface properties of ePTFE facial membranes involved the acrylate monomer monoacryloxyethyl phosphate (MAEP, Fig. 3.6a), using simultaneous grafting in water. Very low graft yields were observed. However, it was clear, from x-ray photoelectron spectroscopy (XPS) and in infra-red spectroscopy investigations, that the phosphate-containing graft copolymer had been successfully introduced (Grøndahl et al., 2002). Subsequent in vitro mineralisation studies using SBF of both phosphate- and carboxylate-modified ePTFE substrates revealed that not HAP but rather the more acidic calcium phosphate phase brushite or monetite formed predominantly, Fig. 3.6a (Grøndahl et al., 2003). The similarities in mineralisation outcomes for these two functionalities were later explained by the complex nature of the MAEP graft copolymer and homopolymer (Suzuki et al., 2006). Indeed it was found that the MAEP monomer is hydrolytically unstable and therefore forms a copolymer of MAEP and acrylic acid. Subsequent studies involved graft polymerisation of the methacrylate monomer methacryloxyethyl phosphate (MOEP, Fig. 3.6b) in various © Woodhead Publishing Limited, 2010

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O O

P

OH OH

P O

O

10 μm (a)

OH OH

20 μm (b)

3.6 (a) MAEP-grafted ePTFE and (b) MOEP-grafted ePTFE showing brushite or monetite (a) and HAP mineral phases (b) respectively.

solvents (e.g. methanol and methyl ethyl ketone (MEK)) (Wentrup-Byrne et al., 2005; Suzuki et al., 2005; Chandler-Temple et al., 2010a). It was found that the graft morphology varied with the solvent used (smooth morphology when grafted in methanol and globular morphology when grafted in MEK) and this was attributed to the solubility of the graft and homopolymers in the respective solvents (Wentrup-Byrne et al., 2005). Although, based on XPS examination, the chemistry of the two graft copolymers appeared to be very similar (i.e. they were both copolymers of MOEP and HEMA), the in vitro mineralisation outcomes were very different (Suzuki et al., 2005). From infra-red spectroscopic analysis of the mineralised layer it was found that only the surface grafted in methanol actually induced nucleation and growth of HAP (Fig. 3.6b). Other modifications resulted in the growth of a mixture of calcium phosphate phases. It was later shown that this difference was in fact due to the degree of crosslinking in the graft-copolymer (Suzuki et al., 2006). Depending on the grafting conditions, the nature of the graft-copolymer varied (Fig. 3.7a or b) and it was shown that a low degree of crosslinking (Fig. 3.7b) was required in order for HAP to nucleate (Fig. 3.6b). In addition, it was found that the crosslinking reactions of the graft copolymer were due to the presence of large amounts of a diene impurity in the monomers used in these studies. Subsequently, a method was devised to produce the non-branched graft copolymer (and homopolymer) without incorporation of the diene impurity (Grøndahl et al., 2008). As a very nice conclusion to this mineralisation study, it was found that the ePTFE membrane containing the non-branched MOEP copolymer induced rapid and pure HAP growth in SBF (Grøndahl et al., 2010). In addition to these promising in vitro mineralisation studies, it has also been shown that the grafted ePTFE membranes produced increased protein © Woodhead Publishing Limited, 2010

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(b)

ePTFE

ePTFE

3.7 Grafted PTFE copolymer showing (a) no crosslinking and (b) some crosslinking.

adsorption and osteoblast attachment, which are both important for increasing the interfacial strength (Suzuki et al., 2005). Our recent work has demonstrated that in vitro macrophage response is affected by the types of proteins that adsorb from serum and can be minimised by judicious choice of monomer and solvent combinations during the grafting process (Chandler-Temple et al., 2010b). In conclusion, through selective modification of the material substrate surface, membranes can be improved (made more bioactive) and this results in stronger interfacing at the facial soft-tissue– bone interface.

3.5

Future trends

It would appear that the most desirable materials are those that are both biostable and capable of integrating at the soft–hard tissue interface. Since structurally graded PTFE materials are already commercially available, this concept is one which lends itself to further exploitation and development. Currently there are plenty of on-going fundamental materials research studies, including the chemical modification of surfaces of several welldefined polymers already used in facial repair and reconstruction. New developments should ideally combine both these aspects with the latest biomedical engineering technologies.

3.6

Acknowledgements

The authors (EW-B, LG and AC-T) would like to thank everyone including their students who, over the years (in particular Drs AC-T and Shuko Suzuki), worked on research projects the results of which are included in this chapter; Dr Richard Lewandowski for his generosity over many years in sharing his expertise in plastic surgery; Drs Bernard Môle and Pierre Nicolau (Paris), both of whom took the time to make a helpful contribution to our discussion on facial atrophy. We thank Sybil Curtis for her drawings in Figs 3.1 and 3.3, Dr Oliver Locos for Fig. 3.4 and finally Patrick J. Lynch,

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medical illustrator and C. Carl Jaffe, MD, cardiologist for permission to use Fig. 3.2 (http://creativecommons.org/licenses/by/2.5/). Thanks are also due to Professor Graeme George whose helpful discussions in the beginning helped to initiate this chapter and who proofread it at the end.

3.7

References

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he c and gu z (2003), ‘Studies on the electron beam irradiated and acrylic acid grafted PET film’, Radiation Physics Chem, 68(5), 873–874. hegazy e-s, dessouki a m, el-assy n b, el-sawy n m and el-ghaffar m a (1992), ‘Radiation-induced graft polymerization of acrylic acid onto fluorinated polymers. 1. Kinetic study on the grafting onto poly(tetrafluoroethylene– ethylene) copolymer’, J Poly Sci Polym Chem, 30, 1969–1976, doi: 10.1002/ pola.1992.080300920. hirsch r j and cohen j l (2006), ‘Soft tissue augmentation’, Cutis, 78(3), 165–172. holck d e and ng j d (2006), Evaluation and Treatment of Orbital Fractures: A Multidisciplinary Approach, Elsevier Saunders, Philadelphia. homicz m r and watson d (2004), ‘Review of injectable materials for soft tissue augmentation’, Facial Plast Surg, 20(1), 21–29. hontsu s, nakamori m, kato n, tabata h, ishii j, matsumoto t and kawai t (1998), ‘Formation of hydroxyapatite thin films on surface-modified polytetrafluoroethylene substrate’, Jpn J Appl Phys, 37, L1169–L1171. hsueh c-l, peng y-j, wang c-c and chen c-y (2003), ‘Bipolar membrane prepared by grafting and plasma polymerization’, J Membrane Sci, 219, 1–13. hürzeler m b, kohal r j, naghshbandi j, mota l f, conradt j, hutmacher d and caffesse r g (1998), ‘Evaluation of a new bioresorbable barrier to facilitate guided bone regeneration around exposed implant threads. An experimental study in the monkey’, Int J Oral Maxillofacial Surg, 27(4), 315–320. isolagen (2007), The Isolagen Process, Isolagen Inc. Accessed: 5 February 2009, . kamei s, tomita n, tamai s, kato k and ikada y (1997), ‘Histologic and mechanical evaluation for bone bonding of polymer surfaces grafted with a phosphatecontaining polymer’, J Biomed Mater Res, 37(3), 384–393, doi: 10.1002/ (SICI)1097-4636(19971205)37:3<384::AID-JBM9>3.0.CO;2-H. kang i-k, choi s-h, shin d-s and yoon s c (2001), ‘Surface modification of polyhydroxyalkanoate films and their interaction with human fibroblasts’, Int J Biological Macromol, 28(3), 205–212, doi: 10.106/S0141-8130(00)00165-3. kannan r y, salacinski h j, butler p e, hamilton g and seifalian a m (2005), ‘Current status of prosthetic bypass grafts: A review’, J Biomed Mater Res, Part B: Appl Biomater, 74B, 570–581, doi: 10.1002/jbm.b.30247. kaufman m r, miller t a, huang c, roostaien j, wasson k l, ashley r k and bradley j p (2007), ‘Autologous fat transfer for facial recontouring: Is there science behind the art?’, Plast Reconstr Surg, 119(7), 2287–2296. kidd k r, dal ponte d b, kellar r s and williams s k (2002), ‘A comparative evaluation of the tissue responses associated with polymeric implants in the rat and mouse’, J Biomed Mater Res, 59(4), 682–689. klein a w and elson m l (2000), ‘The history of substances for soft tissue augmentation’, Dermatol Surg, 26(12), 1096–1105, doi: 10.1046/j.1524-4725.2000.t01-100512.x. kühn g, retzko i, lippitz a and unger w (2001), ‘Homofunctionalized polymer surfaces formed by selective plasma processes’, Surface Coatings Technol, 142– 144, 494–500. lafaurie m, dolivo m, porcher r, rudant j, madelaine i and molina j m (2005), ‘Treatment of facial lipoatrophy with intradermal injections of polylactic acid in HIV-infected patients’, J Acquir Immune Defic Syndr, 38(4), 393–398.

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lam s m, azizzadeh b and graivier m (2006), ‘Injectable poly-L-lactic acid (Sculptra): Technical considerations in soft-tissue contouring’, Plast Reconstr Surg, 118 (3 suppl), 55S–63S. landau t (1989), ‘Identity’, in About Faces, New York, Doubleday, 36–43. lascaratos j g, segas j v, trompoukis c c and assimakopoulos d a (2003), ‘From the roots of rhinology: the reconstruction of nasal injuries by Hippocrates’, Ann Otol Rhinol Laryngol, 112(2), 159–162. lawrence e l and turner i g (2005), ‘Materials for urinary catheters: A review of their history and development in the UK’, Medical Eng Physics, 27(6), 443–453, doi: 10.1016/j.medengphy.2004.12.013. lemperle g, romano j j and busso m (2003), ‘Soft tissue augmentation with Artecoll: Ten-year history, indications, techniques, and complications’, Dermatol Surg, 29(6), 573–587, doi: 10.1046/j.1524-4725.2003.29140.x. linde a, alberius p, dahlin c, bjurstam k and sundin y (1993), ‘Osteopromotion: A soft-tissue exclusion principle using a membrane for bone healing and bone neogenesis’, J Periodontol, 64(11), 1116–1128. lorenzoni m, pertl c, polansky r and wegscheider w (1999), ‘Guided bone regeneration with barrier membranes – a clinical and radiographic follow-up study after 24 months’, Clin Oral Implants Res, 10(1), 16–23. lowe n j, maxwell c a and patnaik r (2005), ‘Adverse reactions to dermal fillers: Review’, Dermatol Surg, 31(s4), 1616–1625, doi: 10.2310/6350.2005.31251. malaisrie s c, malekzadeh s and biedlingmaier j f (1998), ‘In vivo analysis of bacterial biofilm formation on facial plastic bioimplants’, Laryngoscope, 108(11/1), 1733–1738. mazzei r o, smolko e, torres a, tadey d, rocco c, gizzi l and strangis s (2002), ‘Radiation grafting studies of acrylic acid onto cellulose triacetate membranes’, Radiation Physics Chem, 64, 149–160. mazzei r, tadey d, smolko e and rocco c (2003), ‘Radiation grafting of different monomers onto PP foils irradiated with a 25 MeV proton beam’, Nuclear Instrum Methods Physics Res B, 208, 411–415. mercandetti m, cohen, a j and chang e w (2008), Implants, Soft Tissue, Gore-Tex, last updated 8 Jan 2008. Accessed 3 Feb 2009, . morhenn v b, lemperle g and gallo r l (2002), ‘Phagocytosis of different particulate dermal filler substances by human macrophages and skin cells’, Dermatol Surg, 28(6), 484–490, doi: 10.1046/j.1524-4725.2002.01273.x. moyle g j, lysakova l, brown s, sibtain n, healy j, priest c, mandalia s and barton s e (2004), ‘A randomized open-label study of immediate versus delayed polylactic acid injections for the cosmetic management of facial lipoatrophy in persons with HIV infection’, HIV Med, 5(2), 82–87, doi:10.1111/j.1468-1293.2004.00190.x. moyle g j, brown s, lysakova l and barton s e (2006), ‘Long-term safety and efficacy of poly-L-lactic acid in the treatment of HIV-related facial lipoatrophy’, HIV Med, 7(3), 181–185, doi: 10.1111/j.1468-1293.2006.00342.x. neel h b (1983), ‘Implants of Gore-Tex’, Arch Otolaryngol, 109(7), 427–433. neuber g (1893), ‘Fettransplantation’, Zentrabl Chir, 22, 66. nicolau p j (2008), ‘The different products and their pharmacology’, in Ascher B, Landau M and Rossi B (eds.) Injection Treatments in Cosmetic Surgery, London & New York, Informa HealthCare, 413–417.

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stewart m g (2005), Head, Face and Neck Trauma: Comprehensive Management, New York, Thieme. suzuki s, grøndahl l, leavesley d and wentrup-byrne e (2005), ‘In vitro bioactivity of MOEP grafted ePTFE membranes for craniofacial applications’, Biomaterials, 26, 5303–5312. suzuki s, whittaker m r, grøndahl l, monteiro m j and wentrup-byrne e (2006) ‘Synthesis of soluble phosphate polymers by RAFT and their in vitro mineralization’, Biomacromolecules, 7(11), 3178–3187, doi: 10.1021/bm060583q. tanahashi m and matsuda t (1997), ‘Surface functional group dependence on apatite formation on self-assembled monolayers in a simulated body fluid’, J Biomed Mater Res, 34, 305–315. tretinnikov o n and ikada y (1997), ‘Hydrogen bonding and wettability of surface grafted organophosphate polymer’, Macromolecules, 30, 1086–1090. truswell w h (2002), ‘Dural-porosity expanded polytetrafluoroethylene soft tissue implant’, Arch Facial Plast Surg, 4, 92–97. ut southwestern (2009), Auricular Reconstruction (Ear Replacement Surgery), last updated 2009. Accessed: 5 January 2010, . valantin m a, aubron-olivier c, ghosn j, laglenne e, pauchard m, schoen h, bousquet r, katz p, costagliola d and katlama c (2003), ‘Polylactic acid implants (New-Fill®) to correct facial lipoatrophy in HIV-infected patients: Results of the open-label study VEGA’, AIDS, 17(17), 2471–2477. van helden a (1995), ‘Tycho Brahe’, The Galileo Project, last updated 1995. Accessed: 22 September 2008, . vert m, li s m, spenlehauer g and guerin p (1992), ‘Bioresorbability and biocompatibility of aliphatic polyesters’, J Mater Sci Mater Medicine, 3(6), 432–446, doi: 10.1007/BF00701240. voorhees a b j, jaretzki iii a and blakemore a h (1952), ‘The use of tubes constructed from Vinyon “N” cloth in bridging arterial defects’, Ann Surg, 135(2), 332–336. walsh w r (2006), Repair and Regeneration of Ligaments, Tendons and Joint Capsule, Totowa, NJ, Humana Press. wentrup-byrne e, grøndahl l and suzuki s (2005), ‘Methacryloxyethyl phosphategrafted expanded polytetrafluoroethylene membranes for biomedical applications’, Polym Int, 54, 1581–1588. wilkins r h (1964), ‘Neurosurgical Classic – XVII’, J Neurosurg, 21, 240–244. wiltfang j, merten h-a and peters j-h (1998), ‘Comparative study of guided bone regeneration using absorbable and permanent barrier membranes: A histologic report’, Int J Oral Maxillofacial Implants, 13, 416–421. woerle b, hanke c w and sattler g (2004), ‘Poly-L-lactic acid: A temporary filler for soft tissue augmentation’, J Drugs Dermatol, 3(4), 385–389. yamaguchi s, yabutsuke t, hibino m and yao t (2008), ‘Generation of hydroxyapatite patterns by electrophoretic deposition’, J Mater Sci Mater Med, 19, 1419–1424, doi: 10.1007/s10856-006-0053-6. zhao s j, pinholt e m, madsen j e and donath k (2000), ‘Histological evaluation of different biodegradable and non-biodegradable membranes implanted subcutaneously in rats’, J Cranio-Maxillofacial Surg, 28(2), 116–122.

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4 Corneal tissue engineering Y. - X. H UA N G, Ji Nan University, China

Abstract: This chapter reviews the current position in the field of corneal tissue engineering. First, the construction and biophysical properties of the human cornea and the special conditions required for corneal tissue regeneration are illustrated. Then the strategies and approaches used for engineering corneal tissue are introduced, including the traditional way of engineering corneal equivalents by seeding cells in biodegradable scaffolds in vitro; the method of using active cell-free artificial corneas to induce corneal tissue regeneration in vivo; and the methods based on stem cells and cell sheet technology. Key words: corneal tissue engineering, corneal equivalent, active artificial cornea, stem cell, cell sheet technology.

4.1

Introduction

There are several reasons why the cornea is an ideal candidate for tissue engineering research. First, corneal disease is the second most common cause of blindness and there are over 10 million patients worldwide who remain blind from corneal disease through a lack of donors for corneal transplantation. Therefore, there is a strong desire to produce artificial bioactive corneal tissue substitutes to overcome the problem of corneal tissue failure. Second, its construction is less complex than most tissues so the reconstruction of engineered substitutes should be simpler. Third, since it is an avascular tissue with immune privilege, there should be fewer problems regenerating its tissue in vivo. Therefore numerous attempts have been made to engineer corneal tissue using different approaches. Significant progress has been made and engineered corneal epitheliums are already in clinical use. This chapter will review the current position in the field of corneal tissue engineering, and introduce the strategies and approaches used. To enable readers to gain a better understanding of the tissue properties and the specific requirements of corneal tissue engineering, we will also describe the construction of the human cornea; its biophysical properties, especially its optical properties which account for its particular function as a transparent window; and the special conditions for tissue regeneration. Finally, we will discuss future challenges and prospects for corneal tissue engineering. 86 © Woodhead Publishing Limited, 2010

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Characteristics of the human cornea and its regeneration

4.2.1 Construction of the human cornea The cornea is the outermost layer of the eye and represents 7% of a human eye’s surface area. It forms a firm shell with the sclera to protect the components of the eye from infection and injury and also serves as a transparent window for external images to enter the eye. The shape of a human cornea is a concave–convex lens with a radius of curvature of about 7.7 mm at the front surface, but with a value of about 6.8 mm at the back. It is ∼0.5 mm thick at the center, increasing to about 0.7 mm at the periphery. For most people, the corneal curvature on the front surface is different in the vertical and horizontal directions. The horizontal meridian is usually more curved in young corneas, but this reverses as corneas age. When viewed from the front, the cornea has a diameter of 11 mm in the vertical direction, which is slightly smaller than its diameter of 12 mm in the horizontal direction. With its curved shape, the cornea acts like a lens to focus images on the retina at the back of the eye, but causes a certain amount of astigmatism owing to its toricity. Essentially, the cornea is a connective tissue containing collagen in the form of fibrils and it is an avascular tissue. However, it can obtain its physiological requirements from the lacrimal fluid in front of it and the aqueous humor behind it. Structurally, the cornea contains five parallel layers (see Fig. 4.1). The outermost layer is the epithelium, whose role is to absorb nutrients and oxygen while protecting the eye. The epithelium is composed of about five stacks of epithelial cells and has a 4–7 μm thick superficial tear film. The outer epithelium can be divided into three regions: the central cornea, the limbus and the conjunctiva. The avascular cornea is connected to the blood system in the corneal limbus. The blood vessels of the limbus are not responsible for the nutrition of the cornea, but release leucocytes and proteins into the corneal stroma, which plays a role in the diseases and tissue regeneration of the cornea and conjunctiva (Maurice, 1969; Reim, 1982). Below the epithelium is the Bowman’s membrane, which has a thickness of 8–12 μm. The Bowman’s membrane consists of randomly arranged collagen fibrils, mainly Types I, III, V and VII. Beneath it is the stroma, a dense layer of connective tissue making up about 90% of the corneal thickness. The main constituents of the stroma are water (78% of the total weight at normal hydration); collagen (mostly Type I: 15%); glycosaminoglycans (GAGs) (1%); noncollagenous proteins (5%, some of which are bound to the GAGs); and salts (1%) (Fatt and Weissman, 1992). The stroma is composed of a stack of approximately 200 collagenous lamellae, consisting of

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Biointegration of medical implant materials Epithelium Basement membrane

Bowman’s layer

Stroma

Descemet’s membrane

Endothelium

4.1 Construction of the human cornea (adapted from Reichl, 2008).

interwoven collagen fibrils. It contains a population of keratocytes that are derived from neural crest cells, and these occupy 3 to 5% of the total volume (Fatt and Weissman, 1992). The keratocytes are flattened fibroblasts that normally lie quiescently but can be activated in response to injury. They are responsible for the secretion of the unique stromal extracellular matrix, and the synthesis of collagen. Beneath the stroma is a thin limiting lamina called Descemet’s membrane and beneath that, finally, is the endothelium. The endothelium itself is a monolayer of non-dividing cells. The cells are about 20 μm in diameter and 4 to 5 μm thick (Fatt and Weissman, 1992), and they are separated from each other by spaces about 20 nm wide. The intercellular spaces can act as slits to allow different compounds, such as water and oxygen, to diffuse through and thus control corneal hydration, which is crucial for the maintenance of corneal transparency.

4.2.2 Biophysical properties of the human cornea Acting as a firm shell to protect the components within the eye, the cornea has sufficient mechanical strength to maintain excess pressure, which varies between 10 and 21 mm Hg for normal humans (Pinsky et al., 2006), and withstand external knocks and the forces applied by the extraocular muscles during eye movement. The physical properties of the cornea depend on position and direction. Its Young’s modulus was reported (Howland et al.,

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1992; Smolek, 1994; Vito et al., 1989) to be typically about 1–10 MPa, ranging from 2.45 × 104 to 5.7 × 107 Pa, depending on the techniques used for the measurements. It has a high tensile strength along the direction of the collagen fibril axes. Therefore, circumferentially, the cornea is stiffest at the limbus, with a tensile modulus of ∼13 MPa, about 13 times greater than in the horizontal direction (Ruberti et al., 2007). Radially, as viewed from the front, the cornea is stiffest in the central region out to about 4 mm, with maximum strength in the inferior/superior and nasal/temporal directions. In addition, the cornea has very low strength from front to back (i.e. through its thickness), but has high tensile strength along the direction of the collagen fibril axes. As a soft tissue, the cornea is a viscoelastic material. Its strain depends on the time for which a stress acts. There is also a certain amount of creep if the stress is prolonged. Human corneas have a long-term creep component when put under prolonged stress in vitro. As the window of the eye, optically the cornea has several features. Firstly, it is a specialized avascular tissue that maintains a highly organized architecture, but contains no pigments or molecules that would absorb visible light; therefore it is extremely transparent. Its transmittivity is as high as 95% for visible light. At the nanoscopic level, its transparency is achieved from the excellent arrangement of the corneal collagen fibrils within the lamellae. The fibrils are quite uniform in diameter (30.8 nm) (Meek et al., 2003) and are positioned relative to each other with a high degree of lateral order and a relatively constant interfibrillar spacing of 60 nm, which is much smaller than the wavelength of visible light (Fatt and Weissman, 1992). Such an arrangement causes destructive interference of scattered light and constructive interference of directly transmitted light for all the visible wavelengths (Ijiri et al., 2006). Secondly, its refractive index is 1.376. This, together with a convex air– cornea interface, makes the cornea the major focusing component of the eye, acting like a lens to focus incident light on the retina. Over two thirds of the eye’s focusing occurs at its front surface. Thirdly, it is precisely curved and has a smooth surface. Its surface is coated by a tear film that is 4–7 μm thick. Optically, the tear film provides a very smooth surface over the cornea. The precise curvature of the cornea and the optical smoothness of the tear film together make it a highly efficient converging lens. The tear film also transports metabolic products to and from the cornea, preventing the cornea from drying out and thus helping to maintain its transparency and refracting power. Both the transparency and refractive index of the cornea are reported to be critically dependent on the hydration of the tissue (Meek et al., 2003). One more property which should be considered in tissue engineering is the cornea’s permeability. Since the cornea is the main barrier to substances

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entering the eye, the extent to which different compounds permeate the cornea is an important biopharmaceutical parameter. The cornea’s permeability not only represents its barrier property for the penetration profiles of substances, but also influences its hydration. As mentioned previously, the cornea consists of the epithelium, stroma and endothelium, which form the three primary layers through which substances can permeate. For transcorneal transport, the corneal epithelium usually presents a much greater barrier than the stroma. However, the penetration profiles of substances through the corneal tissues depend not only on the cornea’s barrier properties, but also on the chemical nature, size and conformation, lipid/water partition coefficient, and degree of ionization of the permeant molecules (Bijl et al., 1997, 1998, 2000, 2001). Ionic and relatively hydrophilic substances will diffuse slowly through the epithelium, while lipophilic agents will diffuse slowly through the underlying aqueous stroma. The permeability of the human cornea to different substances has been extensively investigated (Fatt and Hedbys, 1970; Weissman et al., 1983; Myung et al., 2006). However, we will mention only results for some of the substances that might be involved in corneal tissue engineering. The water flow conductivity is reported to be 19 × 10−13 cm4 dynes−1 sec−1 at 37 °C with physiological hydration (H = 3.2) (Fatt and Hedbys, 1970). The oxygen permeability is 29.51 × 10−11 ml O2 cm2/sec mL mmHg (SD = 5.62) (Weissman et al., 1983), and the diffusion coefficient of glucose is 3.0 ± 0.2 × 10−6 cm2/s (Myung et al., 2006). It should be noted that the cornea’s permeability strongly depends on its hydration, since a greater water content can decrease the volume fractions occupied by glycosaminoglycans (GAGs) and GAG-associated proteins. An increase in tissue hydration can significantly increase permeability; the accompanying increase in tissue thickness further decreases the tissue’s permeability. Hydration also influences other properties of the cornea. It affects the pressure-induced radial straining of the tissue and hence plays a part in determining corneal shape (Hjordtal, 1995). An increase in tissue hydration can also induce a loss of corneal transparency due to an increase in light scattering from the uptake of water by the stroma. It also induces a change in the corneal refractive index (Meek et al., 2003). Therefore, corneal hydration is a key factor in the biophysical behavior of this tissue.

4.3

Special conditions for wound healing and tissue regeneration of the cornea

In any method being used for corneal tissue engineering, there should be a process for wound healing and tissue regeneration. Therefore, knowledge

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of how the tissue is built, how it grows, how it is repaired, and what kinds of factors/substances are involved in the process, is very useful. We should bear in mind that the cornea is a specialized avascular tissue that must maintain a highly organized architecture in order to maintain its clarity. Therefore, its repair processes include mechanisms that deal with these specialized conditions. Initially, wound healing initiates the ‘reexpression’ of a number of genes; then, as a result, proteins and cellular products are synthesized and interact in a well-orchestrated process (Dayhaw-Barker, 1995a). The following phases usually occur in the process of corneal wound healing and tissue repair: inflammation, epithelialization, fibroplasia, extracellular matrix (ECM) deposition, and remodeling. The substances involved in the process are mesenchymal matrix proteins, proteoglycans, growth factors, proteolytic enzymes, inflammatory mediators and several cell types (Dayhaw-Barker, 1995b). The mesenchymal matrix proteins are responsible for structural, cellular and extracellular adhesion. The main matrix proteins found within the extracellular matrix are: (i) fibronectin (FN), which plays a role in cellular adhesion during migration; (ii) tenascin, a protein that is mainly expressed during either tissue repair or malignant processes; (iii) the integrins, cellular receptors for the extracellular compounds such as FN, laminin (LN), and collagens; (iv) LN, an integral part of most basement membranes, whose role is to promote the adherence of cells to the basement membrane, thus affecting cell movement, differentiation, and growth. Besides these proteins, the most notable fibrillar elements are the collagens, which form a group of at least twelve proteins (Montes et al., 1984; Kuhn, 1987; Burgeson, 1988; Marshall et al., 1993) and can be separated into striated and filamentous collagens. The striated ones are composed of Types I, II, III, and V. The filamentous collagens are primarily Types VI, VII, IX, and X. Each type plays a different role in the function of the cornea, such as increasing tensile strength, serving in some structural capacity, or helping to establish an anchoring site. The dermatan and keratin sulfate proteoglycans are the main proteoglycans or glycosaminoglycans (GAGs) found in corneal tissue. The GAGs not only play a role in maintaining tissue hydration (Katz et al., 1986), but also assist in the spacing and orientation of collagen fibers (Borcherding et al., 1975), which are critical for the tissue’s transparency (Tuft et al., 1993). The growth factors that affect the cornea are: the defensins, which originate from inflammatory cells and have an antimicrobial effect and the ability to stimulate epithelial cell growth (Murphy et al., 1993); the epidermal growth factor (EGF) family; the fibroblast growth factor, which mainly promotes cellular in vitro proliferation and differentiation with possible neurotropic and angiogenic properties; the insulin-like growth factors, such

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as IGFe l, which stimulate mitogenic activity and differentiation; the neuropeptides, which play a role in the stimulation of certain epithelial cell migration patterns (Nishida et al., 1992); the platelet-derived growth factors, which can stimulate the expression of IL-8 mRNA and the fibroblastic production of specific types of collagen and general matrix formation; the transforming growth factors (e.g. TGF-β), which regulate the matrix proteins and their integrin receptors; and lastly the retinoblastoma-derived growth factor, which also plays a role in the stimulation of corneal epithelial cell growth. The enzymes involved in the restructuring of the cornea are the serine proteinases, which include plasmin and its related compounds (Setten et al., 1989; Tervo et al., 1989), and the metal cofactor-requiring enzymes (Fini et al., 1992), including collagenase and the matrix metalloproteinases (MMP). The serine proteinases are responsible for the cleaving of matrix proteins and may influence the orderly progress of cellular migration. The cells involved in corneal tissue regeneration are the epithelial cells, stromal keratocytes and endothelial cells (Dayhaw-Barker, 1995a). The epithelial cells are highly involved in any epithelial insult. The limbus surrounding the cornea is a reservoir of corneal epithelial stem cells. Stromal keratocytes remodel collagen fibers in the implant, heal damaged collagen fibers, and create adhesions in the stromal tissue. The endothelial cells play a key role in transporting ions from the cornea, thus performing the function of a dehydrating pump (Sumide et al., 2006) to maintain the clarity of the cornea. Besides the above-mentioned compounds and cell types in the cornea, it should be noted that there are also some substances needed for wound healing and tissue regeneration in the tear layer, such as the tear proteins lysozyme and lactoferrin, both of which play a role in the defense mechanisms of the cornea and influence its susceptibility to infections (Reim et al., 1997). In addition, since the maintenance of the epithelial cover on the stroma is largely dependent on the wetting of its surface, a tear layer is needed for repairing the epithelial, otherwise dellen and persistent epithelial defects will develop. Various kinds of graft are usually involved in most cases of corneal tissue engineering; the process of tissue regeneration in the grafts is quite different from that of pure wound healing. Therefore, in addition to the information mentioned above, knowledge of the biological response to the grafts, the possibility of graft rejection and the process of re-establishment of normal corneal tissue in the grafts is required. When a graft is implanted, whether it is a corneal graft from a donor, a tissue-engineered corneal graft, or just a synthetic/composite scaffold, successful transplantation and tissue regeneration involve two processes: (i) repopulation of the graft by host cells, especially if the donor cells are

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depleted during the cryopreserve procedure; (ii) the maintenance or regeneration of the graft’s transparency. It has been proved that repopulation of the graft by host cells always happens, whether or not there are donor cells in the graft. The replacement of donor cells by host cells seems to be a normal and gradual process. In the long term, all cell types of the donor tend to become necrotic and be replaced by recipient cells. The epithelium and endothelium are replaced relatively early, within the first post-operative year. However, the replacement of stromal keratocytes takes much longer. It has been reported that a small proportion of stromal keratocytes can even survive for several years but they are all replaced eventually (Wollensak and Green, 1999; Henderson et al., 2001; Binder et al., 1982; Karakami et al., 1991). The repopulation and replacement by host keratocytes can occur without opacification of the graft. The factors that might contribute to a shortened cell cycle or early necrosis of donor cells were suggested to be (Wollensak and Green, 1999; Binder et al., 1982): preoperative cell death in cadaver donor tissue, mechanical trauma during the operation, attrition from aging, the toxic substance from the glycerin-preservation process remaining within the donor tissue, the absence of epithelial stem cells in the transplant, neuroparalytic mechanisms, disturbed metabolism because of the circular scar tissue and the sutures, chronic corneal edema, post-surgical inflammation, elevated intraocular pressure or pseudophakia, and the process of freezing for a cryopreserved graft. A piece of research on rabbits (Karakami et al., 1991) reported that keratocyte regeneration usually occurred within 21 days of implantation. The regenerating ‘keratoblasts’ were characterized by active mitochondria and an increase in the rough endoplasmic reticulum. Keratocytes tended to accumulate along the interface between the host and the corneal graft on day 16 and exhibited marked accumulation on day 45, indicating that keratocytes actively participated in the adhesion and healing of the host tissue and the graft. Following the proliferation and repopulation of keratocytes, collagen synthesis took place. The collagen synthesized by regenerating keratocytes was found to be incorporated more quickly into mature collagen forms. All these results suggest that the regenerated keratocytes play a role in remodeling collagen fibers in the graft, healing damaged collagen fibers in the host stroma, creating adhesions in the host tissue and the graft, and the recovery of corneal clarity. Therefore, to obtain transparency and a suitable refractive index, the donor graft has to be repopulated with the host’s keratocytes, there should be a reestablishment of a normal corneal extracellular matrix, and the stromal collagens should be re-aligned in a regular fashion (Binder et al., 1982, 1986; Zavala and Krumeich, 1987).

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According to the reports of some pieces of research (Ohno et al., 2002; Zhang et al., 2007), the corneal transparency was maintained for just one month after transplantation. In particular, the cryopreserved graft appeared somewhat cloudy throughout the post-operative week. The significant decrease in the transparency of the graft was usually due to a malfunction of any one of the three primary parts of the cornea, such as the loss of epithelium and the decomposition of collagen soon after the operation. In the cryopreserved corneas, most of the keratocytes were killed during the process of freezing, so very few normal-looking keratocytes were visible. On the 10th post-operative day, enlarged, activated keratocytes started to migrate into the periphery of the graft from the host stroma. The fiber diameter and interfibrillar spaces among the collagen fibers became larger than normal with a change in the Bowman’s membrane, which may be one of the reasons why the graft appeared cloudy. Even in the corneal graft which was preserved at just 4 °C before transplantation, the number of keratocytes was reduced to approximately one-third of the number seen in normal corneas that have not been operated on. After this, superficial neovascularization could be observed to grow at the peripheral regions of the graft about two weeks after transplantation and recede gradually from the fourth week. At the same time, the transparency of the graft improved because of the regeneration of the epithelium and the reconstruction of the corneal stroma. By the 16th week after grafting, the uniformly thin collagen fibrils were organized into lamellae, and no significant differences were found between the transparency and neovascularization of the graft and those of normal tissue. A series of remodeling processes will happen after grafting, such as inflammatory cell infiltration, stromal cell proliferation, epithelial mitosis, migration, collagen deposition, and nerve regeneration (Zhang et al., 2007). During the remodeling process, the cells from the corneal graft will become functional by promoting the ingrowth of host cells and the regeneration of damaged nerves from the surrounding tissue. In addition, many cytokines from the stromal cells can regulate the proliferation, motility, differentiation, and possibly other functions of the epithelial cells and eventually accelerate the improvement in corneal clarity (Wilson et al., 1999, 2003). The normal cornea is an immunologically privileged site (Barker and Billingham, 1973; Katami, 1991; Streilein, 1995). There are multiple mechanisms for maintaining its immune privileged status (Niederkorn, 1995, 2002; Niederkorn et al., 1989; Steinman and Nussenzweig, 1980; Tommila et al., 1987), including the lack of blood vessels, lack of lymphatics, blood–eye barrier, relative paucity of mature antigen-presenting cells (APCs) in the central cornea, presence of immunomodulatory factors in the aqueous humor, and the constitutive expression of CD 95 L (Fas ligand) within the

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eye (Panda et al., 2007; Osawa et al., 2004; Prendergast and Easty, 1991; Ray-Keil and Chandler, 1985, 1986; William et al., 1985). This privilege can be destroyed by inflammation and neovascularization. Therefore, steps have to be taken to reduce the risk of inflammation post grafting. Usually, an immunological reaction can be induced only up to six weeks after transplantation, corresponding to the time needed for the replacement of stromal cells. After this period, the donor stromal cells are either no longer present (Maumenee, 1951) or are present only in small amounts (Wollensak and Green, 1999). After an immunological reaction, corneal graft rejection may take place. Corneal graft rejection is defined as a complex immune-mediated process resulting in decompensation of the transplanted cornea. The process is initialized by recognition of the foreign histocompatibility antigens on the cells of the corneal grafts by the host immune system, which results in host sensitization. This is followed by a specific immune response against these antigens, which finally results in the decompensation of the graft tissue. We want to emphasize that corneal graft rejection is primarily a cell-mediated response; exogenetic cells would not only induce immunoreaction, but would also be obstacles to the repopulation of the host’s keratocytes and new tissue generation. Therefore, from this point of view, a graft without donor cells would have a lower risk of graft rejection.

4.4

Approaches to corneal tissue engineering

Recent approaches to corneal tissue engineering can be basically subdivided into four methods, according to their strategies or concepts. The first one, a traditional strategy, aims to seed and culture specific cell types in a biodegradable scaffold with a 3-D structure. The second strategy is based on the use of active scaffolds that closely mimic natural extracellular matrices, which are used in combination with signal biomolecules to induce corneal tissue regeneration in vivo. The third method is purely cell based, involving the dropping/injection of stem-cell suspensions or the transplantation of the cultured cells with supporting materials into a defect site or an injured corneal tissue. The fourth method focuses on the transplantation of autologous or allogeneic cell sheets without using scaffolds.

4.4.1 Traditional: cell-seeded biodegradable scaffold-based tissue engineering The strategy of the traditional method of corneal tissue engineering is to reconstruct a living corneal tissue in vitro by seeding and culturing cells in a biodegradable scaffold with biomolecules that promote the regeneration of tissues. The first step is the isolation and culture of cells. To recon-

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struct complete corneal tissue with all the three major cellular layers mentioned in Section 4.2.1, all the cell types in the corneal tissue (the epithelial cells, stromal keratocytes and endothelial cells) are needed. Since the stem cells located in the limbal portion surrounding the cornea exhibit extensive proliferation potential (self-renewal capacity) and appropriate differentiation abilities, they are usually taken as the best cell source for epithelial cells. The epithelial cells can be isolated from the limbus by dispase digestion using the method described by Green (Rheinwald and Green, 1975) that promotes epithelial cell proliferation. Human corneal keratocytes, on the other hand, can be obtained from the keratocyte contamination in epithelial cell cultures by changing the culture medium for a keratocyte culture medium. Since the medium is not adequate for the proliferation of epithelial cells, only keratocytes will be left in the subcultures. Of course, corneal keratocyte cultures can also be established from the corneal stroma (Germain et al., 2000). Human corneal endothelial cells are usually isolated by the enzymatic treatment of excised corneas (Engelmann et al., 1988). After obtaining the cells, the second step is the production of the stroma tissue. Different kinds of material have been used as the scaffold for reconstructing the stroma tissue, such as human Type I and III collagen (Germain et al., 1999), bovine Type I collagen (Minami et al., 1993), polyglycolic acid (Hu et al., 2004) and a collagen–chondroitin sulfate substrate (Griffith et al., 1999). The reconstructed stromal tissue is produced by mixing corneal keratocytes with the materials, then pouring the mixture into a Petri dish containing an anchorage ring and incubating for gelation. Fibroblasts reorganize the extracellular matrix on the culture into a better substrate for epithelial cells. The stroma reconstructed in this way resembles the corneal stroma but the Descemet’s and Bowman’s membranes are absent (Germain et al., 1999). The next step is to reconstruct the layers of the epithelium and endothelium. Epithelial and endothelial cells are seeded and cultured respectively on the top and the bottom of the reconstructed stroma under either AIC (air-interfaced culture) or LCC (liquid-covered culture) conditions for several more days (Germain et al., 1999; Schneider et al., 1999). The resulting corneas have been found to be histologically similar to a native cornea and express components of the epithelial basement membrane at the epithelium–stroma junction. Although some of the engineered corneal tissues constructed in this way have shown morphology similar to a native cornea, including well-defined epithelial, stromal, and endothelial cell layers, they have been used mainly as models for physiological, toxicological, or pharmacological investigations, as they are not yet ready to be used for transplantation. The reason is that most of them do not have enough tensile strength for surgical manipulation and

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fixation. A detailed study of these kinds of corneal equivalent is still needed to further improve their tensile strength for surgical applications.

4.4.2 Active scaffold which induces corneal tissue regeneration in vivo As mentioned in the last section, corneal equivalents constructed using the traditional method of tissue engineering do not have enough tensile strength to form a surface curvature or to allow surgical manipulation and fixation for transplantation. In addition, in order to construct corneal equivalents, several cell lines (epithelial cells, stromal keratocytes and endothelial cells), at least, have to be prepared and cultured for the regeneration of the different layers of the cornea. The processes of culturing and proliferating each cell type on substrates are complicated and time consuming. Furthermore, even if such a corneal equivalent can be built and has enough tensile strength to be transplanted into a recipient’s eye, we have learned from Section 4.3 that to obtain transparency and a suitable refractive index, the donor graft has to be repopulated with the host’s cells. The previously seeded allogeneic cells in the graft tend to be necrotic and are replaced by recipient cells just after transplantation. They may be acting as a biological medication and stimulate proliferation of the recipient’s residual cells in some way, such as producing extracellular matrix proteins as well as cytokines and growth factors. However, while they induce immunoreaction, the dead donor cells will also be obstacles to the repopulation of the host’s cells. Moreover, in most cases, cell-seeded grafts have to be cryopreserved before transplantation. It has been proved that most of the keratocytes will be killed during the process of freezing; even in a corneal graft which was preserved at just 4 °C before transplantation, the number of keratocytes was reduced to approximately one-third of that seen in normal corneas that have not been operated on (Ohno et al., 2002). Therefore, the cultured donor cells are not actually necessary, since the positive role they may play, if any, comes mainly from what they produce during culturing, such as extracellular matrix proteins, cytokines and growth factors. On the other hand, some reports about the continuous maturation and differentiation of immature or incomplete tissues after transplantation (Compton et al., 1989; Germain et al., 1995) suggest that there is a sophisticated regulating system in the human body that guides and controls the growth of different cells, with each kind of cell being guided to grow in their proper location to form different tissues. Therefore, it would be better to let the human body do the work as a bioreactor to culture cells and generate new tissues in vivo. Based on these considerations, in 2002, Huang et al. proposed a novel strategy for corneal tissue engineering (Huang and Li, 2002a, b, c, 2003,

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2007). They suggested using a cell-free active artificial cornea to induce new corneal tissue formation in vivo. The active artificial cornea they produced had not only a similar composition (such as collagen and extracellular matrix) to the human cornea and the required optical and biomechanical properties, but also some signal molecules such as growth factor. They believed that an acellular scaffold with such properties could induce the recipient’s cells to grow in it then repopulate it and generate new corneal tissue, as happens in an implanted corneal graft. The in vivo ‘bioreactor’ can provide not only the most suitable environmental conditions, but also all the substances needed for tissue regeneration as mentioned in Section 4.3. Since there were no dead donor’s cells that needed to be removed to let the host cells migrate in, the tissue regeneration of the acellular graft was expected to occur in an even shorter period than for a corneal allograft (Huang and Li, 2007). The active artificial cornea was designed to mimic the human cornea and was composed of collagen, chitosan and glycosaminoglycans (GAGs). Collagen is the major structural protein in the human corneal stroma and has many desirable features. It has low immunoreaction and the ability to promote normal tissue regeneration. Chitosan has excellent biocompatibility with the human body. Its interaction with collagen can inhibit collagenase and protease to reduce the biodegradation of the collagen. It also has a high tensile strength to support the formation of collagen into different shapes. The third material, glycosaminoglycan (GAG), was cross-linked to form the extracellular matrix for cell adhesion, migration, proliferation and tissue remodeling. It also gave the new artificial cornea a higher elastic modulus and a more porous structure. Type I collagen powder and chitosan were dissolved in water of pH < 4.5, and the concentration of collagen was adjusted to 1% by acetic acid aqueous solution. The solution was stirred by a homogenizer for 30 min, and then the GAG was slowly added and homogenized for 45 min. The GAG/ collagen–chitosan homogenous solution was spread on a glass plate and heated at 35 °C for 48 h to obtain a completely dry membrane. The dry membrane was then immersed into deionized water for l h to be modified. The artificial cornea displayed similar optical and mechanical properties to a human cornea. It had an optical transmittance of ≥95%; a refractive index of 1.368∼1.382; a tensile strength of 167 kg/cm2 (dry state); an ultimate elongation of 300%; and its degree of swelling in water was 65–75%. With such physical properties, the artificial cornea substrate can easily be constructed into different shapes and curvatures for transplantation (see Fig. 4.2), and will be well tolerated by the recipient. Animal (rabbit) artificial corneal implantation was performed to prove the efficacy of this approach. The implants were prepared in the shape of

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4.2 Artificial cornea lenticules with the function of inducing corneal tissue regeneration in vivo. The diameter of the lenticules is 8.05 mm, the radius of their surface curvature is 6.76 mm.

4.3 Specimen taken on the 60th day. Some host’s keratocytes accumulated along the interface between the host and the graft, and grew into the superficial lamellae of the implant; no inflammatory cell was observed.

discs with a diameter of 8 mm and thickness of 0.3 mm, then rinsed and aseptized in a sterile PBS buffer. In the implantations, the rabbit corneas were excised down to the middle of the stromal layer at the center of the cornea. The active artificial cornea discs, after being administered with a growth factor (bFGF), were inserted into the recipient corneal beds, and two interrupted nylon sutures were used to close the incision. Although the animals were given the injection against anti-immunologic rejection only on the 15th day, the corneas of the animals were observed to maintain a smooth surface and clear stroma as well as transparency. The eyes were clear and had only a little blood-vessel hyperplasia on the 10th day, and no inflammation and hyperemia on the 30th day. A specimen taken on the 60th day, shown in Fig. 4.3, indicated that the implant and the native corneal tissue had good compatibility and connection. Some keratocytes could be seen to accumulate along the interface between the host and the graft and these grew into the superficial lamellae of the implant, while no inflammatory cells were observed. The biodegradability measurement showed

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4.4 Specimen taken on the 120th day post-operatively. Only a small portion of the implant was still degrading, accompanied by metrocytes growing at 4 months. New corneal tissue formation can be observed in the region taken by the implant originally (inside the dashed frame). The newly generated tissue and the host’s deep stroma have the same tinctorial properties; its collagens have good regularity and integration with the native ones.

that there was 30% biodegradation between the 45th and 60th day post-operatively. Figure 4.4 shows a specimen taken on the 120th day post-operatively. It can clearly be seen that the implant was almost totally degraded. In the region occupied by the implant originally (the region in the dashed frame), new corneal tissue had been generated by the repopulating keratocytes, and there were plenty of corneal metrocytes in the zone where the artificial cornea was degrading. The newly generated tissue and the host’s stroma had the same tinctorial properties; its collagens had good regularity and integration with the native ones. The follow-up observation showed that the new corneal tissue generation and implant degradation were completed between the 120th and 150th day. Complete tissue regeneration in the control group using corneal allografts as implants required an extra month. The new generation of the corneal tissue was proved by the specimen taken 6 months after the transplantation. It was about 0.3 mm thicker than the normal (control) cornea. This indicates that it really did contain newly generated cornea tissue. Slit-lamp microscope imaging on the eye of the animal just before sacrifice 5 months post-operatively showed that the

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cornea was clear, smooth and homogeneous in thickness. Animal experiments of lamellar keratoplasty using the novel artificial cornea also showed similar results, and complete re-epithelialization on the surface of the cornea with the implant was found on the 5th day after transplantation. In later experiments performed in a similar way by other authors, using materials such as hydrated collagen and N-isopropylacrylamide copolymerbased ECMs to fabricate the implants, tissue regeneration was also observed in the acellular implants in rabbits and mini-pigs (Griffith et al., 2008; Liu et al., 2006). The efficacy of this approach was further proved. In summary, the strategy of using an active acellular scaffold to induce host cell repopulation and new cornea tissue generation in vivo opens up a new approach to corneal tissue engineering. Since there was no exogenetic cell, the active scaffold had only a small rejection reaction, mainly induced by the suture after implantation. On the other hand, its GAG and growth factor components had an anti-inflammatory effect and gave the autologous cells the signal to grow, so new tissue was generated faster in the active acellular scaffold than in the one using homologuous corneal lamina for the implants. As we mentioned previously, corneal transplantation has a special requirement that other tissue transplantation does not: to obtain transparency and a suitable refractive index, the donor implant has to be repopulated with the host’s keratocytes, and the stromal collagens should be re-aligned in regularity (Binder et al., 1982, 1986; Zavala and Krumeich, 1987). So no matter whether it is a homologuous cornea or a cornea equivalent produced in a typical tissue engineering process, they both have to undergo such a process. They have no advantage over an artificial cornea without cell culturing before transplantation. On the contrary, their exogenetic cells may induce immunoreaction and would be obstacles to the repopulation of the host’s keratocytes and new tissue generation. Therefore, a cell-free artificial cornea that induces new corneal tissue generation in vivo would be a better method for corneal transplantation if the host can still generate keratocytes. Its advantage is fairly obvious: not only can it be used universally for different implantation surgeries for various people, but it is also suitable for long-term storage and longdistance transportation. In addition, by having similar optical and mechanical properties to the human cornea, the novel artificial cornea is well tolerated by the patient and is easy to prepare and process into different shapes and sizes on a large scale. It can also be processed into lenticules with different optical powers for epikeratophakia. Its immediate applications include increasing the thickness of a cornea for patients with a thinner cornea or critical corneal thickness, resulting from radial keratoplasty or other causes such as wearing contact lenses. It can also be used for refractive keratophakia.

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4.4.3 Stem-cell tissue engineering It is well known that stem cells have the unique capability to self-renew and to generate committed progenitors. They can differentiate into the cell lineages of the tissue of origin after a limited number of cell divisions, so stem cells have been widely used for engineering different tissues, including corneal tissue. For corneas, the self-renewal of corneal epithelium can usually be achieved by the stem cells located in the limbal region of the cornea. However, some pathological ocular conditions, such as severe chemical or thermal burns, genetic disorders, microbial keratitis, collagen vascular diseases and immunological disorders, can cause permanent damage of these stem cells. When limbal stem cells are destroyed, the conjunctival epithelial cells may migrate to cover the denuded corneal surface. Such a resurfacing would result in an irregular opaque ocular surface as well as stromal neovascularization, chronic inflammation, and persistent epithelial defects. Conjunctivalization of the cornea may lead to irreversible corneal scarring and impaired vision. The traditional way to repair the corneal epithelium and restore vision in these patients was to perform limbal tissue transplantation or combined limbal corneal transplantation, but this was restricted by a shortage of donor corneas. Therefore, a method of engineering corneal tissue from allogeneic or autogenetic (in case of unilateral limbal stem-cell deficiency) stem cells was developed to overcome this difficulty (Lehrer et al., 1998; Sangwan et al., 2003; Schwab et al., 2000b; Tsai et al., 2000; Sun and Green, 1977). Compared with whole organ transplantation, this method has several advantages: it can provide a potentially unlimited resource for corneal epithelial stem-cell transplantation; it is an easy, safe and less painful procedure; and it can offer the potential for in vitro manipulation of the graft prior to implantation to eliminate unwanted features such as the depletion of antigen-presenting cells from cultivated cell populations. The clinical use of cultured limbal stem cells was first reported by Pellegrini et al. in 1997 (Pellegrini and Traverso, 1997). Since then, many reports of the clinical use of this technology have been published, and in several clinical centers it has become a routine stem-cell therapy for the permanent regeneration of a corneal epithelium in patients with limbal stem-cell deficiency. According to Pellegrini, epithelial cell cultures (Pellegrini and Traverso, 1997) were established from 1–2 mm2 full-thickness biopsy samples taken from the ocular surface (bulbar conjunctiva, cornea, and limbus) of donors. The biopsy samples were minced and treated with trypsin at 37 °C for 3 h. The cells were then plated on a feeder layer of lethally irradiated 3T3-J2 cells (2.4 × 104/cm2) and cultured in 5% carbon dioxide in Dulbecco-Vogt

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Eagle’s and Ham’s F12 media (3:1 mixture). Before transplanting the cultured limbal epithelial cells, the conjunctival epithelium covering the cornea and limbus of the injured eye had to be removed. After placing the cultured epithelial graft on the prepared eye, a soft therapeutic hydrophilic contact lens was placed over the graft to provide an efficient protective bandage to shield the graft and provide a structural framework upon which the regenerating epithelium could grow. The transplantation of limbal epithelial cells cultured from an organ donor has to be performed in combination with immunosuppressive therapy, and the approach has a very low success rate in the long term, as compared with autologous cells. Interestingly, it was found that the majority of successful cases eventually showed only recipient DNA in the regenerated epithelium (Henderson et al., 2001). This is not so surprising, since we have mentioned in Section 4.3 that over the long-term in tissue regeneration, all cell types of the donor tend to become necrotic and be replaced by recipient cells, though this time, the recipient’s limbal stem cells were believed to be destroyed, and some of the cases even had total limbal stem-cell deficiency. Therefore, it was suggested that allogeneic cultures were acting as a biological medication and actually stimulating the proliferation of residual stem cells of the patient (Pellegrini et al., 2007). Owing to the fragility of the epithelium, the clinical application of cultured human limbal cells in this way could not be reproduced on a large scale (Pellegrini and Traverso, 1997). Therefore, some supporting materials have been used for cell culture, transportation and transplantation onto patients, such as fibrin glue (Rama, 2001), amniotic membrane (Koizumi, 2001; Sangwan et al., 2003; Schwab et al., 2000; Tsai et al., 2000), collagen sponges or strips, and devitalized membranes or polymers (Nishida et al., 2004b). Although providing the cultured epithelium with a basement membrane is likely to improve graft ‘take’, and may even promote the survival of cultured stem cells (Daniels et al., 2001b), alternative culture methods have, nevertheless, been tried. Many efforts have been focused on the removal of the feeder layer and/or fetal calf serum, on the assumption that there are potential risks from the use of these supporting materials. Another idea being considered is that using new culture media or cultivating limbal cells onto different carriers, even with cells in suspension, may help the transplantation to be performed by the dropping or injection of stem-cell suspensions into a defect site or an injured corneal tissue. In one of the approaches not using a feeder-cell layer (Qian et al., 2001), the cells were cultured in the following way: freshly excised limbal tissue was immersed in Dispase II (1.2 unit/mL) for 4 hours at 37 °C under 5% CO2. Epithelial stem cells were collected by scraping them off the limbal tissue. The cell suspension was then centrifuged at 1500 rpm for 10 min and

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washed in MEM once. The cell pellet was re-suspended in the culture medium and seeded in a 36 mm culture dish. Cells were cultured at 37 °C under 5% CO2 for 2 days with a culture medium. The medium was then changed to a serum-free keratinocyte culture medium for another 10 to 14 days. When the growth of epithelial cells reached a confluent layer (1–3 × 105) in 10 to 15 days from one corneal strip, cells were released by trypsin and transplanted on the 14th day after limbectomy. An average of 3 × 105 cells in 0.2 mL of limbal cell culture medium were transplanted in the form of eye drops. Cells were transferred to both the conjunctival sac and a soft therapeutic hydrophilic contact lens, which was then placed to cover the ocular surface. No significant graft rejection was found with allogeneic stem-cell transplantation, even in the absence of immunosuppressive therapy. The method seems promising, since corneal epithelium was regenerated in 30 out of 36 (83.3%) allograft and 35 out of 41 (85.3%) autograft eyes. However, it has been used only on rabbits and its clinical performance on humans has not yet been proved. Until now, the main aims of stem-cell therapy have been to promote the re-epithelialization of the cornea, provide a stable epithelium, prevent regression of new vessels, and restore epithelial clarity. Can similar methods be applied to the stroma and endothelium of the cornea? Does the presence of a graft acting as a biological dressing reduce the hostility of the local environment such that any remaining host stem cells can survive and function? Why does a transplanted cell migrate and differentiate appropriately or not, and how does the environment stimulate such parameters? These are questions which require further investigation. Human embryonic stem cells (hESCs) could be a potential source of corneal cells for generating corneal tissues for transplantation. The advantages of using hESCs are obvious, such as the possibility of studying the pathways involved in cell-lineage differentiation and having an unlimited source of corneal epithelial cells for transplantation. Committed stem cells have been obtained from mouse embryonic stem cells for corneal reconstruction in mice (Ahmad et al., 2007) and corneal epithelial cells have also been produced from hESCs (Homma et al., 2004). It has even been proposed that stem cells could be transported to diverse tissue compartments via the blood stream, and that on arrival the cells would generate the appropriate cell types in response to local signals (Daniels et al., 2001a).

4.4.4 Corneal tissue engineering based on cell sheet technology Although the potential for stem-cell tissue engineering is significant, there are still some critical problems. For instance, with the dropping/injection of single-cell suspensions, in many cases the implanted cells cannot be retained

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around the target tissue, so it is difficult to control their size, shape and location after the transplantation. The use of supporting materials makes handling the cell culture easier, but their presence can potentially influence the post-operative clinical outcomes. For example, an amniotic membrane would persist between the corneal stroma and the expanded epithelial cells after transplantation (Kinoshita et al., 2004), and could even affect the optical transparency of the fabricated constructs. Although fibrin gel degrades rapidly after transplantation into the corneal stroma, its biodegradation can induce inflammation and thus might possibly result in microtrauma to the corneal stroma. Moreover, the possibility of infection from the use of biological carriers cannot be completely excluded (Yang et al., 2006). To solve these problems, an alternative method called cell sheet technology has been developed, which involves the creation of carrier-free constructs that can still be easily manipulated during surgical operations (Yamada et al., 1990; Okano et al., 1993, 1995). Cell sheet technology mainly employs a ‘thermo-responsive culture dish’ to enable reversible cell adhesion to and detachment from the dish surface by the controllable hydrophobicity of the surface (Yamada et al., 1990; Okano et al., 1993, 1995). The technology enables a non-invasive harvest of cultured cells to be made as an intact monolayer cell sheet, including the deposited extra-cellular matrix (ECM). The monolayer cell sheet can be collected simply by reducing the culture temperature to below 32 °C for less than 30 minutes, without using any enzymes or chelating agents. The technology also enables the direct transplantation of cell sheets into host tissues without scaffolds, fixation, or sutures. The main material used for the thermo-responsive graft-cell culture dish is poly(N-isopropylacrylamide) (PIPAAm). The temperature-responsive polymer exhibits thermo-responsive hydrophobicity changes in aqueous solutions (Heskins et al., 1968). By means of electron beam irradiation of an IPAAm monomer on polystyrene dishes, the monomer is polymerized and covalently bonded with the dish surface. At temperatures below 32 °C, PIPAAm molecules are highly hydrated, so the PIPAAm grafted surfaces are hydrophilic. However, when the temperature rises above 32 °C, the surfaces suddenly change to hydrophobic due to the extensive dehydration of the PIPAAm molecules. When cells attach to a material surface, they physically adhere onto hydrophobic surfaces better than onto hydrophilic surfaces. After a period of passive adhesion, cells change their morphology to spread due to metabolic processes using ATP. On the hydrophobic surface of a thermoresponsive culture dish, adhered cells can proliferate normally to confluency and express normal phenotypic markers for each cell type. Spread cells rarely change their shape to a round one in order to detach from the hydrophobic surface without the release of surface-engaged receptors. When the

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Suspended cells

Cell adhesive at 37 °C

Corneal epithelial cell sheet

Donut-shaped PVDF membrane

Non-adhesive at 20 °C

Temperature reduction

Transplanted to ablated stromal bed

After removal of donut-shaped supporter

4.5 The procedure of epithelial cell sheet preparation and transplantation (Hayashida et al., 2006). PVDF, polyvinylidene difluoride.

temperature is below 32 °C (typically 20 °C for 30 minutes), the surface becomes hydrophilic, and spread and adhered cells can be spontaneously detached from the thermo-responsive culture dish surface due to hydration of the surface. This detachment process does not damage the cells because the temperature change is a physically mild treatment, and no enzymes such as trypsin or a chelator such as EDTA are required. The monolayers of cells have even been proved to maintain basal surface ECM proteins after detachment (Kushida et al., 1999). It has been proved that various kinds of cell can be harvested on the PIPAAm grafted culture dish under general incubation conditions of 5% CO2 and 37 °C (Matsuda et al., 2007). Figure 4.5 shows the procedure for epithelial cell sheet preparation and transplantation (Hayashida et al., 2006). By using cell sheet technology, one can reconstruct tissues in three general ways. For tissues such as the corneal epithelium, single cell sheets can be directly transplanted into host tissues (Nishida et al., 2004a). To create 3-D structures, one can perform homotypic layering of cell sheets (Shimizu et al., 2002, 2003). For more complex laminar structures, the method of stratifying different cell sheets should be considered (Harimoto et al., 2002).

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Cell sheet technology has been successfully applied to corneal tissue engineering (Nishida et al., 2004a). In the reconstruction of the corneal epithelium, limbal epithelial stem cells were isolated and cultured on the temperature-responsive culture surfaces at 37 °C. After the cells reached confluency, the culture temperature was decreased to 20 °C for 30 min, when the cell sheets, together with their deposited ECM, could be easily harvested and adhered to the host corneal stroma without the need for sutures (Nishida et al., 2004a). Since the corneal epithelial cell sheets were harvested without using dispase, they were less fragile and contained both cell-to-cell junctions and ECM proteins that can be damaged by dispase. It was demonstrated that a well-formed epithelial sheet can be transplanted without the need for any carrier substrate, such as an amniotic membrane or fibrin gel, and the transplanted cell sheets can attach to the corneal stroma in 5 minutes without the need for sutures. It was found that in patients receiving corneal epithelial cell sheet transplantation, the corneal surface remained clear, with significantly improved visual acuity, more than one year after surgery. In cases of severe disease or bilateral limbal stem-cell deficiency, damage to both eyes prevents the use of autologous corneal epithelial stem cells. In these circumstances, Nishida et al. demonstrated that autologous oral mucosal epithelial cell sheets could be used as an alternative to corneal epithelial cell sheets (Nishida et al., 2004b, Hayashida et al., 2005). They found that oral mucosal epithelial cell sheets fabricated on temperatureresponsive dishes could be harvested and transplanted in the same manner as corneal epithelial sheets. The oral mucosal epithelial sheets fabricated in this way resembled the native corneal epithelium even more closely than the native oral mucosal epithelium. Results from all the human trials they performed demonstrated remarkably improved visual acuity, with all corneas maintaining a clear and smooth surface. They also found that in a rabbit model, after transplantation, oral mucosal epithelial cell sheets underwent changes in their keratin expression profiles towards a corneal phenotype (Hayashida et al., 2005). They believed that the results indicated that the direct interaction between the transplanted cell sheets and the underlying corneal stroma might have had an effect on the phenotypic modulation of the oral mucosal epithelial cells. These findings are encouraging, since using autologous oral mucosal epithelial cell sheets to replace allogeneic corneal epithelial cell sheets can prevent many of the problems induced by allogeneic immunorejection and immunosuppression. They also allow patients with bilateral limbal stem-cell deficiency to be treated easily and provide a sufficient source of epithelial cell sheets for transplantation. However, the interpretation of the findings remains somewhat unclear, and there are questions that need answering, requiring further investigations.

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It was also demonstrated that corneal endothelial cell sheets (Nakamura and Kinoshita, 2003; Nishida et al., 2004b; Sumide et al., 2006) can be created in a similar way. The results of transplanting human corneal endothelial cell (HCEC) sheets into a rabbit model showed that the implanted HCEC sheets were able to maintain a viable monolayer structure beneath the corneal stroma that closely resembled the native endothelium. In summary, cell sheet technology has the following advantages. As the cultured cells are harvested only by a mild, low-temperature treatment without the use of proteolytic enzymes, cell sheet technology allows the non-invasive harvest of cultured cells as an intact monolayer cell sheet including deposited ECM (Kushida et al., 1999). By means of their deposited ECM, cell sheets can be directly attached to host tissues and even wound sites without scaffolds, fixation, or sutures, with minimal cell loss and fewer inflammatory responses.

4.5

Future trends

Over a development period of almost 20 years, although corneal tissue engineering has made significant progress and shown great promise for supplying engineered bioactive organ and tissue substitutes, it has not advanced as fast as people expected. As yet, it is still a great challenge to fabricate a complete substitute for the whole corneal tissue with all the three primary layers, and mimic its optical and mechanical properties. The problems and obstacles are numerous. For instance, although the cornea is a less complex tissue than most of the tissues in the human body, at the same time it has special properties, since it requires high transparency and a specific refractive index, besides possessing specific biomechanical properties. Even though it consists of only five layers, it is composed of many substances. In the extracellular matrix of its stoma alone, there are eight types of collagens and some other substances such as GAGs. Each of them plays a certain role in the biophysical properties of the cornea. They are organized in regularity to ensure transparency, tensile strength, and hydration, as mentioned in Section 4.3. However, in constructing a corneal substitute using the present techniques of tissue engineering, it is very difficult to fabricate one with all the substances that appear in the human corneal stroma in similar proportions, although the techniques for building 3-D corneal constructs composed of all the three cell types of human cells already exist. More difficult is the establishment of their particular organization, which accounts for the transparency, tensile strength and hydration. In addition, unlike the in vivo ‘bioreactor’, the in vitro bioreactor at present cannot provide all the environmental conditions and substances needed for corneal tissue regeneration to culture a corneal substitute with the structure and properties that mimic the human cornea. Furthermore, after a graft is

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implanted, the repopulation of host cells and the remodeling of collagens always occur in order to maintain or re-establish the transparency of the implanted graft. Therefore, an engineered cell-seeded graft, especially the allogeneic cell-seeded ones, will face more problems of immunological reaction and inflammatory responses, because corneal graft rejection is primarily a cell-mediated response. The greater problem is whether its integration could be supported by the host tissue. Therefore, the strategy and approaches for further development of corneal tissue engineering should be reconsidered. Do we need to construct a complete substitute for the whole corneal tissue or just replace its damaged portions? In most cases, the cornea is only partly damaged. So on the whole, only substitutes for the damage portions are required; a complete corneal substitute with all the three primary layers is seldom needed. For this particular circumstance, at present, cell sheet technology is probably the right way to supply engineered epithelium and endothelium. The technique of inducing tissue regeneration in vivo by an active cell-free artificial cornea, on the other hand, is the most suitable means of achieving regeneration of the stroma. A combination of the two techniques, therefore, is expected to be able to regenerate the whole cornea. Certainly stem-cell technology, especially embryonic stem-cell technology, may provide a more efficient means of achieving regeneration of the whole cornea in the future. As embryonic stem-cell culture and differentiation technologies improve, it is hoped that committed stem cells will be obtained to reconstruct the whole cornea in vivo/vitro, or stem cells will be transported to diverse tissue compartments of the cornea via the bloodstream or other means, and that on arrival, the cells can generate the appropriate cell types and tissues in response to local signals. Besides these issues, there are two more puzzles: should we use just autologous cells (including autologous oral mucosal epithelial cells) to culture engineered corneal tissue or also use allogeneic cells? Should the tissue be cultured in vitro or in the body? We hope that this chapter will help readers to discover the right answers.

4.6

References

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5 Tissue engineering for small-diameter vascular grafts J. I. R O T M A N S, Leiden University Medical Centre, The Netherlands; and J. H. C A M P B E L L, University of Queensland, Australia

Abstract: This chapter discusses the current status of research in the field of vascular tissue engineering. The chapter first reviews the clinical need for tissue engineered blood vessels and the required characteristics of engineered vascular constructs. Subsequently, a variety of approaches for vascular tissue engineering are discussed that have been utilized in an effort to design the optimal arterial substitute. Finally, the major hurdles on the way to widespread clinical use are reviewed. Key words: vascular tissue engineering, prosthetic vascular graft, scaffold, stem cells.

5.1

Introduction

5.1.1 The global burden of vascular disease Atherosclerotic vascular diseases, in the form of coronary artery and peripheral vascular disease, are the leading cause of mortality in the western world (Lopez et al., 2006). Although the incidence of cardiovascular events seems to have decreased slightly in the last ten years, a renewed increase is expected in the coming decades due to the global epidemic of obesity and diabetes mellitus (King et al., 1998). Many patients require surgical procedures for replacement of diseased blood vessels in case of critical ischemia. In the United States, more than 1.4 million arterial bypass operations are performed each year (DeFrances and Podgornik, 2006). In addition, 1.6 million patients on chronic hemodialysis worldwide require frequent surgical interventions to create and maintain adequate vascular access (Schwab et al., 1999). Indeed, surgical interventions relating to hemodialysis access are becoming one of the most common procedures performed by a typical vascular surgeon. Unfortunately, adequate tissue for vascular grafting is lacking in many patients due to previous surgical interventions and preexisting vascular disease. Between 10 and 40% of patients do not have veins suitable for grafting owing to pre-existing vascular disease, vein stripping, or vein harvesting for prior vascular procedures (Piccone, 1987). 116 © Woodhead Publishing Limited, 2010

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5.1.2 Limitations of prosthetic grafts In the absence of suitable autologous tissues, non-absorbable prosthetic grafts are frequently used. Today, the two most common prosthetic graft materials are polyethylene terephthalate, or Dacron and expanded polytetrafluorethylene (ePTFE). These prosthetic grafts do not match the efficacy of native vessels, particularly in small-diameter applications. Dacron and ePTFE grafts are much stiffer than the elastic arteries to which they are attached, and lack an anti-thrombogenic endothelial layer at the luminal side (Sarkar et al., 2006). When used as bypass arteries that are less than 6 mm in diameter, thrombosis rates are greater than 40% after six months (Sayers et al., 1998). Similar failure rates are observed when prosthetic grafts are used as vascular access for hemodialysis (Schwab et al., 1999). The limited durability of prosthetic grafts relates to (i) acute thrombosis caused by the absence of functional endothelium, (ii) formation of stenotic lesions due to proliferation of vascular smooth muscle cells (i.e. intimal hyperplasia) and (iii) graft infection (Rotmans et al., 2005b).

5.1.3 Clinical perspective of vascular tissue engineering In view of the above limitations, there is a considerable clinical need for alternatives to autologous veins and prosthetic grafts used in surgical procedures such as lower limb arterial bypass grafting, arteriovenous grafts for hemodialysis and coronary artery bypass grafting. Tissue engineering, defined as the combination of cells, scaffolding, and signaling to form biologically active tissues (Riha et al., 2005), offers potential alternatives to fulfil this urgent clinical need. In fact, tissue engineered blood vessels (TEBV) could ultimately be superior to venous transplants since they would provide diameter-matched conduits without existing disease, valves and bifurcations. Furthermore, TEBVs may have the potential to grow and adapt to changing hemodynamic circumstances, when implanted in pediatric patients. As a consequence of the enormous burden of cardiovascular disease, the development of highly sophisticated biomedical technologies, as well as the increasing knowledge about (stem) cell biology, vascular tissue engineering is a rapidly emerging field in medical research. Since the first attempt by Weinberg and Bell (1986) to engineer new blood vessels, techniques to construct adequate TEBVs have witnessed substantial improvements. However, none of the currently available TEBVs fulfil the high-demanding requirements for arterial substitutes. In the current chapter, we describe several methods that have been utilized in an effort to produce artificial arteries. Furthermore, we will address the main limitations of the current TEBVs and attempt to sketch the main hurdles that need to be taken before successful clinical application becomes reality.

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5.2

Required characteristics of tissue engineered blood vessels

5.2.1 Anatomy and function of native arteries In the field of tissue engineering, it is commonly thought that the best way to successfully mimic the mechanical behavior of native arteries is to engineer tissue with similar composition and structure. Therefore, a successful approach in vascular tissue engineering starts with sufficient knowledge about anatomy and functions of these tissues. Native blood vessels have a concentric layered structure, with each layer being distinct in its cell and protein composition (Fig. 5.1). The tunica intima is the inner layer of the vessel wall that consists of a monolayer of specialized endothelial cells (ECs) together with a subendothelial layer which is composed of connective tissue. This layer possesses a variety of functions, including prevention of clot formation and inflammation of the underlying tissue, as well as signaling functions to the muscular component of the vessel wall (Cines et al., 1998). The internal elastic lamina (IEL) consists of crosslinked elastic fibers and forms the partition between the tunica intima and the tunica media. Elastin is the dominant extracellular matrix protein deposited in the arterial wall and can contribute up to 50% of its dry

V

A

M I

E

EC

Layer I Tunica intima Endothelial cells Glycocalyx Subendothelial extracellular matrix M Tunica media Smooth muscle cells Collagen fibers Elastic fibers Glycoproteins

Function Prevents coagulation Prevents inflammation Signaling to muscle layer

Vasomotion Tensile stiffness Tensile elasticity

A Tunica adventitia Fibroblasts Collagen fibers Vaso vasorum

Nutritional supply Structural support Role in vascular repair

E Elastic laminae Elastic fibers

Tensile elasticity

5.1 Schematic representation of the arterial wall and its function. The left panel is a combined figure of an elastin van Gieson-stained section (left half) and a schematic cartoon of the arterial wall (right half). V, vaso vasorum; EC, endothelial cell.

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weight (Karnik et al., 2003). The tunica media is the muscular layer of the artery which is formed by contractile vascular smooth muscle cells (VSMCs), elastin, and collagens Type I and III, as well as proteoglycans. The mechanical properties of native arteries rely on all these components of the tunica media (Bank et al., 1996). The collagen fibers are aligned circumferentially along the axis of the vessel and provide tensile stiffness (Gelse et al., 2003), required for resistance against rupture. Proteoglycans such as heparan sulfate, chondroitin sulfate, and dermatan sulfate contribute to the compressibility of the arterial wall (Gandley et al., 1997). Elastin is arranged as extracellular concentric cylinders in the native artery, confers elasticity of the vascular wall, and acts as a potent autocrine regulator of VSMC proliferation (Patel et al., 2006). In response to stimuli from the ECs or cytokines from the blood, VSMCs contract or dilate in a coordinated fashion which, in turn, leads to constriction or dilation of the artery. In addition, innervation of the medial layer of larger arteries procures direct communication between the nervous system and the vasculature. Peripheral to the tunica media lies another layer of elastic fibers called the external elastic lamina (EEL) which separates the medial layer from the tunica adventitia. The latter layer consists of a loose collagen matrix with embedded fibroblasts. Small blood vessels called vaso vasorum are also present in this layer, which provides nutritional supply for the vessel wall. Besides provision of structural support, the adventitia is thought to play a role in vascular repair after injury (Li et al., 2000).

5.2.2 Required qualities of tissue engineered blood vessels Vascular tissue engineering aims to obviate the twin obstacles of thrombogenicity and immunogenicity by using autologous cells and a suitable matrix with ideal mechanical properties. Proposed criteria for adequate mechanical properties of TEBV include a burst pressure of >1700 mmHg (L’Heureux et al., 2007a) and suture holding strength of >50 grams-force (Konig et al., 2008). Furthermore, TEBV should retain axial and radial compliance, thereby preserving resistance against aneurysm formation, even up to several years after implantation. In addition, TEBVs should be nontoxic, noncarcinogenic, leak-proof and, in the case of hemodialysis access grafts, resistant to repetitive punctioning. Finally, they should not be susceptible to infection and allow manufacturing in a relatively short space of time with differing specifications in terms of diameter and length. Unarguably, these extensive requirements need to be assessed in long-term follow-up studies in relevant animal models before clinical application in humans can be considered. In order to reach this ultimate goal, close collaboration of vascular biologists with specific expertise in cell-matrix interactions, (chemical) engineers and clinicians is of utmost importance.

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5.3

Approaches to vascular tissue engineering

Tissue engineering has been described as ‘an interdisciplinary field that applies the principles and methods of engineering and the life sciences toward the development of biological substitutes that restore, maintain and improve tissue function’ (Fuchs et al., 2001). In the field of vascular tissue engineering, a variety of methods have been utilized to synthesize new vascular conduits including (i) cell seeding of conventional non-absorbable prosthetic grafts, (ii) in vitro engineering of TEBVs using scaffolds based on synthetic polymers or natural macromolecules, (iii) use of decellularized allografts or xenografts as scaffolds and (iv) the use of self-assembled scaffolds (Table 5.1). In scaffold-based approaches, the scaffold functions to bring the various cells of the vasculature in close proximity to one another to enable intercellular signaling, which in turn facilitates cell adhesion, migration and differentiation.

5.3.1 Seeding of non-absorbable prosthetic grafts Prosthetic grafts have been used in vascular surgery since 1952. Shortly after their introduction, researchers concluded that the patency of smalldiameter prosthetic grafts was limited due to thrombosis, intimal hyperplasia and infection (Kapadia et al., 2008). Since then, several attempts have been made to modify graft surface in an effort to reduce thrombogenicity, to decrease the development of intimal hyperplasia and to improve host incorporation and healing.

Cell seeding of prosthetic grafts The absence of endothelial layer and contact of blood components to the prosthetic material are thought to contribute to graft failure. In humans, prosthetic graft endothelialization is very slow and usually does not extend beyond 1–2 cm of the graft edges (Rahlf et al., 1986). Consequently, ECseeding at the luminal surface of prosthetic vascular grafts is a valuable strategy to improve graft patency. In an effort to restore a physiological endothelial barrier, Herring et al. (1978) were the first to show the benefits of seeding ECs on polyethylene prosthesis in a canine model. Subsequent clinical trials in humans using ePTFE grafts seeded with venous ECs for infra-inguinal bypasses, reported increased patency rates for up to nine years of follow-up (65% for the endothelialized group versus 16% patency of non-seeded grafts, p = 0.02) (Deutsch et al., 1999). Microscopic analysis

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Scaffold

ePTFE

ePTFE

ePTFE

ePTFE Polyurethane

ePTFE

ePTFE

ePTFE Collagen

Fibrin

Elastin

First author (year)

Deutsch (1999)

Bhattacharya (2000)

Griese (2003)

Rotmans (2005) Batchelor (2003)

Dorrucci (2008)

Cagiannos (2005)

Lee (2007) Weiberg (1986)

Swartz (2005)

Leach (2005)

VSMCs

Paclitaxel ECs/VSMCs/ fibroblasts ECs/VSMCs

Sirolimus

Heparin

Circulating EPCs NO-donor

Circulating EPCs

BM-EPCs

ECs

Cells/compound

N/A Burst strength, 323 mmHg 7% of arterial break tension Elastic modulus, 900 mmHg

N/A

N/A

N/A N/A

N/A

N/A

N/A

Mechanical properties

–

Sheep

Pig –

Pig

Human

Pig Sheep

Rabbit

Dog

Human

Animal/human

All grafts (n = 2) patent at 15 weeks Not tested in vivo

65% patency at 9 years of follow up Accelerated endothelialization Accelerated endothelialization 3-fold increase in IH Increased thrombus-free surface 2-year primary patency rate of 85% Decreased stenosis in the outflow graft 100% patency at 12 weeks Not tested in vivo

Results from in vivo studies

Table 5.1 Summary of evaluated strategies for vascular tissue engineering. Abbreviations: ePTFE, expanded polytetrafluorethylene; ECs, endothelial cells; BM-EPCs, bone marrow endothelial progenitor cells; VSMCs, vascular smooth muscle cells; N/A, not available; IH, intimal hyperplasia; PLGA, poly[lactic-co-(glycolic acid)]; PLA, polylactic acid; PCL, polyεcaprolactone; Decell., decellularized

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Hyaluronan

Polyglycolic acid

PLGA

PCL–PLA

PCL–PLA

Decel. porcine aorta Decel. porcine vein Decel. bovine ureters Decel. intestinal mucosa Self-assembled in vitro Self-assembled in vivo Self-assembled in vivo

Arrigoni (2006)

Nklason (1999)

Iwai (2004)

Shin’oka (2001)

Matsumura (2003)

Amiel (2006)

Cheu (2004)

Sparks (1973)

L’Heureux (2006)

Shell (2005)

Chemla (2008)

Katzman (2005)

Scaffold

First author (year)

Table 5.1 Continued

Tissue capsule

Tissue capsule

ECs/fibroblasts

None

None

None

ECs/VSMCs/ fibroblasts Bone marrow cells ECs

ECs/VSMCs

ECs/VSMCs

VSMCs

Cells/compound

Burst strength >2500 mmHg

Burst strength >2000 mmHg N/A

N/A

N/A

Burst strength >1000 mmHg N/A

N/A

33% of arterial tensile strength 24% of arterial burst strength 900% of arterial tensile trength N/A

Mechanical properties

Dog

Human

Human

Pig

Human

Human

–

Human

Human

Dog

Pig

–

Animal/human

73% patency at 6.5 months

80% patency at 5 months (ongoing study) 20% patency at 6 months

Similar patency as ePTFE grafts Similar patency as ePTFE grafts No significant reduction in IH

Patent at 3 years in pulmonary circulation No complications in all 17 patients Not tested in vivo

100% patency at 6 months

100% patency at 24 days

Not tested in vivo

Results from in vivo studies

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of nine failed seeded-grafts at a mean follow up of 39 months confirmed the presence of an endothelium on all specimens (Deutsch et al., 2008). In these clinical studies, a two-stage procedure was utilized in which harvested ECs were expanded in vitro for four weeks before the seeded graft was implanted in the same patient. As a consequence, this time- and costintensive method is not suitable for emergency surgery. To circumvent these limitations, one-stage procedures have been developed in which cell harvest, seeding and graft implantation can be completed during one surgical procedure. Unfortunately, none of the clinical studies using a one-stage procedure showed favorable results over non-seeded prosthetic grafts (Bordenave et al., 2005). As an alternative to mature ECs, recent studies have focused on the use of endothelial progenitor cells (EPCs) for cell seeding procedures. EPCs are a subset of CD34+ cells with the potential to proliferate and differentiate into mature ECs (Asahara et al., 1997). Several studies have emphasized that circulating EPCs have the capacity to home to sites of vascular injury, thus promoting the process of re-endothelialization (Walter et al., 2002; Werner et al., 2003). Initial seeding studies were performed by Bhattacharya et al. (2000), who harvested CD34+ EPC from the bone-marrow of dogs. Subsequently, these cells were seeded on polyester vascular grafts using a two-stage procedure. At four weeks after graft implantation in the thoracic aorta, the authors observed a significant increase in graft endothelialization (92% versus 26.6% in non-seeded grafts, p < 0.01). The feasibility of EPC-seeding in a two-stage procedure was confirmed by Griese et al. (2003), who harvested CD34+ EPC from peripheral blood of rabbits. Again, accelerated endothelialization was observed at four weeks after grafting. Unfortunately, the authors did not report on the effect of EPC-seeding on the degree of neointimal formation in the anastomotic area. In an effort to ‘auto-endothelialize’ prosthetic grafts, we evaluated the efficacy of anti-CD34 coated ePTFE grafts (Rotmans et al., 2005a). Their ability to capture EPCs in vivo was assessed in a porcine model of arteriovenous graft failure in which prosthetic grafts were implanted between the carotid artery and the jugular vein (Rotmans et al., 2003). In spite of cellular coverage of the luminal surface, we observed a three-fold increase in neo-intima formation in anti-CD34 coated grafts after 28 days of follow-up. Obviously, the attracted cells enhanced proliferation of VSMCs in the anastomotic area or transdifferentiated into VSMCs themselves. Additional studies will reveal whether and to what extent these captured cells can be stimulated to regain their endothelial and vasculoprotective function. In recent years, several other cell populations, including CD14+ monocytes (Harraz et al., 2001), mesenchymal stem cells (Oswald et al., 2004) and VSMCs (Wang et al., 2006) have been shown to possess the ability to differentiate into functional

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ECs. Future studies will show if these cells are a suitable source for cell seeding procedures. Graft seeding with small molecules In view of the important role of nitric oxide (NO) in maintaining vascular homeostasis, NO-releasing systems have been developed for stent and graft coatings. NO is a potent vasodilator, with protective effects on ECs. Furthermore, NO inhibits thrombus formation and proliferation of VSMCs (Vallance and Collier, 1994). Since NO has a half-life of only a few seconds, several researchers have focused on the immobilization of NO-donor compounds within biomaterials. Although different studies using NO-releasing grafts showed favorable effects on thrombus formation and intimal hyperplasia (Smith et al., 1996; Batchelor et al., 2003), concerns have been raised regarding the formation of toxic metabolites and the finite reservoir of NO that exists within prosthetic grafts (Reynolds et al., 2004; Kapadia et al., 2008). Currently, clinical data on the use of NO-donor coated grafts are lacking. Protein coating or binding proteins to grafts is another approach for improving graft patency. Recently, several commercially available and FDA-approved grafts have used heparin-coating technologies to improve graft patency (Kapadia et al., 2008). Although no prospective, randomizedcontrolled clinical trials have been conducted, the two-year primary patency rate of 85% of heparin-bonded ePTFE grafts used for infragenicular arterial bypass surgery suggests superior patency rates, when compared to historical controls (Dorrucci et al., 2008). Although these studies are promising, the risk of heparin-induced thrombocytopenia, which can be lethal, remains a drawback (Kapadia et al., 2008). Alternative compounds for graft seeding include sirolimus and paclitaxel. Indeed, sirolimus- and paclitaxel-eluting stents have been used successfully in occlusive coronary artery disease, and have been shown to reduce in-stent restenosis (Eisenberg and Konnyu, 2006). Both these agents possess strong anti-proliferative effects on VSMCs. Lee et al. (2007) assessed the potential of paclitaxel-eluting prosthetic grafts, which were evaluated in a porcine arteriovenous graft model. When compared to conventional prosthetic grafts, paclitaxel-eluting grafts showed better survival than uncoated grafts at 12 weeks after implantation (100% versus 25%, respectively, p = 0.01). Cagiannos et al. (2005) used sirolimus-eluting grafts as iliac artery bypass grafts in pigs and showed a significant reduction in per cent cross-sectional narrowing as well. Future clinical trials will reveal if these encouraging results translate into increased patency rates of prosthetic grafts in patients who require an arterial bypass or hemodialysis access.

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5.3.2 In vitro tissue engineering using scaffolds of natural macromolecules The archetypal tissue engineering approach is to use a scaffold that is seeded with vascular cells in vitro. The properties of an ideal scaffold have been identified as (i) sufficient porosity to allow cell/tissue growth as well as transport of nutrients, (ii) suitable surface chemistry for cell attachment and proliferation and (iii) pseudo-physiological mechanical properties (Hutmacher, 2001). Natural macromolecules such as collagen as scaffold components have the advantage of facilitating cell binding. Collagen scaffolds Collagen Type I is often used in this context because it is abundant in many tissues and can be isolated, solubilized and subsequently poured on a mold of the desired (tubular) shape. Weinberg and Bell (1986) published the first study in which a cell-seeded collagen scaffold was used to engineer a TEBV in vitro. In their study, a multilayered tube was constructed that included bovine ECs, VSMCs, as well as fibroblasts, mimicking the structure of native arteries. Their initial constructs were so highly distensible that they ruptured at very low pressures (<10 mmHg). Tissue strength was improved by integrating a multilayered Dacron mesh sleeve into the collagen lattice. By doing so, a burst strength of 323 ± 31 mmHg was achieved which is still much lower than the recommended burst strength of 1700 mmHg that is required for safe application of TEBV in clinical practice (L’Heureux et al., 2007a). The lack of strength is most likely related to the absence of elastin, the longitudinal (instead of circumferential) orientation of collagen fibers, and the low density of VSMCs. In subsequent studies, increased tissue strength was achieved by collagen crosslinking with glutaraldehyde (Gratzer et al., 1996). However, this approach showed detrimental effects on cell-viability (Huang-Lee et al., 1990), making it unsuitable for clinical application. Alternative methods to induce collagen crosslinking included transfection of VSMCs to overexpress lysyl oxidase, the enzyme responsible for crosslinking of collagen in most tissues (Elbjeirami et al., 2003). However, none of the methods have resulted in collagen hydrogels of sufficient strength and compliance for implantation into the human arterial system thus far. Fibrin scaffolds Fibrin is a blood clotting protein and is involved in wound healing. Although fibrin is not an ECM protein in native arteries, it is a promising scaffold for TEBVs. Indeed, fibrin stimulates synthesis of collagen (Clark et al., 1995) and elastin (Long and Tranquillo, 2003) in VSMCs and fibroblasts, thereby

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improving mechanical properties of TEBVs. Furthermore, incorporation of growth factors in fibrin matrix is facilitated by the natural affinity of certain growth factors for fibrin (Sakiyama et al., 1999). Swartz et al. (2005) engineered implantable small-diameter blood vessels based on ovine smooth muscle and endothelial cells embedded in fibrin gels. These constructs were implanted into the external jugular veins of sheep. Fifteen weeks after surgery, explanted TEBVs exhibited an endothelial monolayer and multiple layers of VSMCs that were circumferentially oriented and produced substantial amounts of collagen matrix and elastic fibers. Concurrently, vasoreactivity and tensile strength appeared to be comparable with native jugular veins. Most importantly, the TEBV remained patent and demonstrated blood flow comparable to unoperated jugular veins. Thus far, these fibrin-gel based TEBVs have not been implanted in humans. Elastin scaffolds In view of the crucial role of elastin in vascular homeostasis and biomechanical function, many investigators turned their attention towards elastinbased scaffolds. However, elastin is highly insoluble and thus essentially unprocessable (Daamen et al., 2001). In an effort to circumvent this limitation, Leach et al. (2005) synthesized elastin-based scaffolds by processing elastin with the hydrophilic crosslinker ethylene glycol diglycidyl ether. Next, the resultant water-soluble crosslinked elastin was used to synthesize thin films on which VSMCs were cultured. In vitro, these cells appeared to adhere relatively well and showed slower proliferation rates when compared to polystyrene tissue culture plastic. To date, in vivo experiments using these scaffolds are lacking. Another method for creation of elastin scaffold has been by Buttafoco et al. (2006) who prepared meshes of solubilized elastin and/or collagen by means of electrospinning. This method has been widely studied to prepare scaffolds with high porosity and surface area, suitable for tissue engineering of small-diameter blood vessels (Vasita and Katti, 2006). Again, VSMCs were successfully cultured on these scaffolds but in vivo data are lacking. The crucial characteristic of these constructs relates to the enzymatic degradation of elastin-based scaffolds by elastase-type endopeptidases, which has been observed in previous studies (Patel et al., 2006). Therefore, scaffold modifications to ensure continuous elastin synthesis or prevention of elastin degradation are essential for the future clinical success of elastin-based scaffolds. Hyaluronan scaffolds Hyaluronan is a non-sulfated glycosaminoglycan which is present in skin, connective tissue, human serum, the vasculature and many other organs

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(Laurent et al., 1996). Several researchers have attempted to create hyaluronan-based scaffolds for vascular tissue engineering (Arrigoni et al., 2006). Indeed, hyaluronan is an important component of the glycocalyx, the network of polysaccharides on top of the endothelium which is crucial for endothelial preservation in peripheral arteries. However, clinical applicability of the current hyaluronan-based scaffolds is limited by their rapid degradation rate, weak attachment of cells and poor mechanical properties (Couet et al., 2007).

5.3.3 In vitro tissue engineering using scaffolds of synthetic macromolecules The major limitations of natural macromolecules are their inappropriate mechanical properties (compared to those of arteries) and the long time required for the construct to mature with cells. Therefore, synthetic macromolecules are designed to obtain tissue engineered constructs with greater mechanical integrity in a (relatively) short time. Synthetic scaffolds have the advantages of industrial replication and reproducible mechanical qualities. Since the scaffold needs only to provide temporal mechanical support to the engineered vascular tissue, biodegradable and biocompatible polymers are the most suitable source of scaffolds for vascular tissue engineering. Careful selection of an appropriate scaffold is of vital importance since the phenotype of seeded cells and the composition of newly expressed matrix proteins are strongly influenced by the scaffold on which they are seeded (Kim et al., 1999). Most of the synthetic polymers used in this field involve materials that are already in widespread use in surgery, such as in sutures. Polyglycolic acid-based scaffolds Niklason et al. (1999) used biodegradable poly(glycolic acid) (PGA) scaffolds that were seeded with VSMCs. Subsequently, these constructs were placed in an in vitro pulsative perfusion chamber for 8 weeks. Compared to engineered vessels grown under static condition, pulsative perfusion enhanced myosin heavy chain expression in VSMCs, a late marker of VSMC differentiation. In addition, pulsative perfusion strongly enhanced collagen synthesis. Next, these vessels were seeded with autologous ECs whereupon they were grafted into the saphenous artery of Yucatan pigs. After 24 days of follow up, all four implanted grafts remained patent with no signs of stenosis or thrombosis. In contrast, non-pulsed control grafts thrombosed at 3 weeks after implantation, which may relate to the loss of endothelial cells due to sudden exposure to arterial flow and pressure. Additional studies from the same laboratory focused

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on mechanistic properties of these tissue engineered grafts (Dahl et al., 2007). Although collagen density matched that of native arteries, burst pressure (803 mmHg) and compliance (3.5% per 100 mmHg) were much lower than those of native arteries (3323 mmHg and 18.7% per 100 mmHg, respectively). To explain this observation, the authors suggest that collagen alignment rather than collagen density determines the mechanical properties of vascular tissues. Indeed, their TEBVs had less circumferentially aligned collagen fibers than native vessels. In addition, the lack of elastic fibers contributed to the low compliance of the engineered vessels as well. In subsequent studies, these researchers designed human TEBVs using a PGA-based scaffold seeded with human VSMCs and ECs (Poh et al., 2005). Unfortunately, these TEBVs also lacked sufficient burst strength to allow successful implantation. Furthermore, they observed that residual polymer fragments and PGA degradation products deteriorate tissue strength and induce phenotypic modulation of VMSCs from a contractile to synthetic state (Higgins et al., 2003). The latter is an important characteristic of intimal hyperplasia and subsequent formation of occlusive lesions in prosthetic grafts. Polylactic acid-based scaffolds Polylactic acid (PLA) is a biodegradable thermoplastic polyester that can be produced through ring-opening polymerization of lactic acid. PLA is a methylated version of PGA, but it is less hydrophilic and therefore degrades slowly (Couet et al., 2007). Mooney et al. (1994) described the fabrication of the copolymer poly[lactic-co-(glycolic acid)] (PLGA) made of PLA and PGA. In subsequent in vivo studies, Iwai et al. (2004) evaluated PLGA–collagen scaffolds seeded with VSMCs and ECs, and implanted them as a patch for the pulmonary artery trunk in dogs. All grafts remained patent up to six months after implantation while histological analysis showed a structure close to that of native arteries, including elastin and collagen fibers. To our knowledge, PLGA-based scaffolds have not been implanted in the human vasculature. Polye-caprolactone-based scaffolds Polyε-caprolactone (PCL) is a biodegradable polyester with remarkable elasticity, which allows it to closely match physiological values. Shin’oka et al. (2001) produced a PCL-based vascular construct and used it for reconstruction of pulmonary outflow tract in a child with atresia of the pulmonary artery. For this purpose, they harvested autologous ECs, VSMCs

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and fibroblasts from a segment of peripheral vein and cultured them for 8 weeks. The scaffold for these cells was composed of a PCL–PLA copolymer (weight ratio, 1 : 1) reinforced with woven PGA. At three years of follow-up, the patient was doing well, with no evidence of graft occlusion or aneurysm formation (Matsumura et al., 2003). Although larger clinical trials are required, tissue engineered vascular grafts might be the ‘Holy Grail’ for children with congenital defects. Indeed, the currently used prosthetic vascular grafts lack growth potential and therefore require replacement as the children grow. TEBVs however, may have the potential to grow and adapt to changing hemodynamic circumstances. Furthermore, the low blood pressure in the pulmonary circulation may favor successful application of TEBVs since the biomechanical properties required from TEBVs in the pulmonary vascular bed are not as highly demanding as those for the systemic arterial circulation.

5.3.4 Decellularized scaffolds Theoretically, transplantation of allogeneic blood vessels is another method to circumvent the shortage of suitable autologous grafts required for surgical procedures. However, immunoreactivity and rejection due to antigenic differences between the donor tissue and the recipient hampers its applicability. Therefore, attempts have been made to remove cells from tissuederived scaffolds using chemical and mechanical treatments (Gilbert et al., 2006). Subsequent addition of autologous cellular constituents can be performed in vitro (Kaushal et al., 2001), preferably after the seeded cells have been cultured under pulsatile flow conditions. Indeed, such method has been shown to increase cell retention after implantation in the arterial circulation (Dardik et al., 1999). Alternatively, the scaffold can be implanted directly into the recipient without prior cell seeding. The latter approach premises spontaneous cell repopulation in vivo, a phenomenon that has been observed in animal studies (Lantz et al., 1992; Clarke et al., 2001). The sources of these biological scaffolds can be divided into three categories: (i) xenogeneic blood vessels obtained from animals (Kaushal et al., 2001), (ii) non-vascular xenogeneic conduits such as intestinal mucosa (Lantz et al., 1992) and ureters (Clarke et al., 2001) and (iii) allogeneic blood vessels obtained from human donors (Schaner et al., 2004). Decellularized scaffolds have the advantage of possessing an ECM that contains cell signaling components such as basic fibroblast growth factor (bFGF) and vascular endothelial growth factor (VEGF) (Voytik-Harbin et al., 1997), which are essential for cell adhesion, proliferation and differentiation (Yow et al., 2006). Preserved ECM components of decellularized blood vessels also include collagen, elastin and glycosaminoglycans. As a consequence, reten-

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tion of tensile strength is ensured (Schaner et al., 2004). Furthermore, these ECM components have the capacity to inhibit clot formation and proliferation of VSMCs (Koyama et al., 1996; Bingley et al., 1998; Li et al., 1998). Another potential advantage of vascular substitutes based on decellularized scaffolds is their greater resistance to infection compared with synthetic grafts (Jernigan et al., 2004). The latter may relate to rapid development of vaso vasorum in biological scaffolds, which in turn facilitates recruitment of leucocytes (Lantz et al., 1993). Xenogeneic decellularized vascular scaffolds Amiel et al. (2006) studied decellularized porcine aortic scaffolds that were seeded with human ECs obtained from the saphenous vein. Analysis of these TEBVs showed intact matrix architecture and a burst pressure of more than 1000 mmHg. Up to now, these decellularized aortic conduits have not been tested in vivo. However, bovine mesenteric veins have been utilized as bioprostheses for hemodialysis access in humans (Katzman et al., 2005). After decellularization using glutaraldehyde crosslinking, these scaffolds were implanted in 183 patients on chronic hemodialysis who required a new vascular access. Unfortunately, the one-year primary patency rate (defined as any event that caused loss of graft patency or required an intervention to the lumen of the graft) did not differ significantly from a controlgroup of 93 patients in which an ePTFE graft was implanted (36% versus 28%, respectively, p = 0.52). Xenogeneic decellularized non-vascular scaffolds A different approach has been executed by Matsuura and co-workers. They assessed the efficacy of decellularized bovine ureters (SynerGraft®) which were implanted without prior cell seeding as arteriovenous grafts in dogs (Matsuura et al., 2004). After 12 months of follow-up, these SynerGrafts® did not show superior primary patency rates when compared to ePTFE grafts (59% versus 57%, respectively). Recently, the results of a randomized clinical trial were published in which these Synergrafts® were compared with ePTFE-grafts in 56 hemodialysis patients without suitable options for creation of a native arteriovenous fistula (Chemla and Morsy, 2008). Again, primary patency rates at one year of follow-up were not better than those of prosthetic grafts (28% versus 48%, respectively, p = 0.29). The investigators observed an excess of early thrombotic events in the SynerGraft® group. This may relate to residual xenoantigens (Spark et al., 2008) or to the absence of a functional endothelium. Indeed, direct contact of platelets with collagen initiates the formation of a hemostatic plug (Nieswandt and Watson, 2003).

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Another material that has been used in this context for more than two decades is small intestinal submucosa (Lantz et al., 1992). This matrix is usually of porcine origin and consists of the fibrous protein-rich layer with collagen and proteoglycans as its main components. In vivo studies in pigs showed superior infection resistance of these scaffolds when compared to ePTFE grafts. However, no significant difference in neointimal response was observed (Shell et al., 2005). Human data on these scaffolds as vascular substitutes are lacking. In summary, there are currently no clinical data available that support the widespread use of decellularized xenogeneic scaffolds as vascular grafts. Furthermore, concerns over the use of xenogeneic scaffolds have been raised which relate to transmission of animal retroviruses and cross-species infection (Yow et al., 2006). Although decellularization reduces the risk of contamination, recent studies have shown that up to 2% of donor DNA remains detectable on some scaffolds after decellularization (Kallenbach et al., 2004). Nevertheless, this is an attractive method for creating an artificial artery because it does not require a graft maturation time, can be used ‘offthe-shelf’, and the construct materials are in abundant supply. Allogeneic decellularized vascular scaffolds Recently, Teebken et al. (2009) reported on successful preclinical development of tissue-engineered vascular substitutes using re-endothelialized human vein matrix. Intact collagen and elastin networks, as well as complete acellularity, was shown after decellularization using sodium deoxycholic acid and DNase I-treatment. Subsequently, ECs obtained from saphenous vein segments were seeded on these scaffolds under different flow conditions. Unfortunately, no data on mechanical properties of these TEBV were included in their report. Thus far, the efficacy of these allogeneic vascular conduits has not been assessed in clinical trials.

5.3.5 Vascular tissue engineering of exclusively autologous blood vessels All the above-mentioned scaffolds for vascular tissue engineering possess limitations that may hamper clinical applicability. Scaffolds based on natural macromolecules usually lack sufficient tensile strength while cytotoxicity of degradation products from synthetic macromolecules has raised concerns about their applicability. Besides antigenicity, potential limitations of decellularized scaffolds relate to their dense collagen and elastic fiber networks, which may hamper cell infiltration and subsequent repopulation (Lu et al., 2004). These drawbacks have resulted in several attempts to create selfassembled TEBVs.

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In vitro engineering of self-assembled vascular grafts More than 10 years ago, L’Heureux et al. (1998) reported a novel approach for vascular tissue engineering based exclusively on the use of cultured human cells, i.e. without any synthetic or exogenous biomaterials. In short, a cohesive sheet made of cultured human VSMCs was placed around a temporal tubular support to produce the media of the vessel. A similar sheet of fibroblasts was wrapped around the media to provide an adventitial layer. After 8 weeks of maturation, the tubular support was removed and ECs were seeded on the luminal surface. These TEBVs displayed excellent mechanical properties with a burst strength of >2000 mmHg. In subsequent studies, solely fibroblasts and ECs were used (L’Heureux et al., 2006). When implanted as interposition grafts in the abdominal aorta of nude rats, 86% of these TEBVs remained patent up to 32 weeks of follow-up. Furthermore, no signs of luminal narrowing or aneurysm formation were observed. Subsequently, autologous TEBVs were implanted as arteriovenous access grafts in six patients on chronic hemodialysis (L’Heureux et al., 2007b). Total production time for the graft ranged between 6 and 9 months and graft length varied from 14 to 30 cm. In the first patient, the TEBV was used for more than 13 months, until the patient underwent successful kidney transplantation. One patient died of unrelated causes while thrombotic failure was observed in one patient at 12 weeks, which was attributed to low flow rate. In the remaining three patients, the vessel was still patent at 5 months after implantation. Furthermore, time to hemostasis after dialysis access was lower for the TEBVs than for ePTFE grafts (L’Heureux et al., 2007a). Recently, the authors reported 60% primary patency rate at 6 months after implantation (McAllister et al., 2009). These preliminary results are encouraging, especially since arteriovenous hemodialysis accesses represent the most challenging mechanical environment in the field of vascular surgery. Besides the repetitive punctioning of the conduit for dialysis, the latter relates to a turbulent flow pattern at the anastomotic areas and the high flow rate which usually exceed 800 mL/min. The major limitations of this in vitro approach are the laborious and time-consuming procedures (28 weeks) (L’Heureux et al., 2006) that are required to grow vital and sterile blood vessels.

In vivo engineering of self-assembled vascular grafts The ultimate goal of these in vivo approaches is that patients grow their own arteries within their own body. The rationale for these methods stems from the observation that implantation of prosthetic materials in the human body initiates an inflammatory response that culminates in the formation of an autologous fibrocellular capsule. Proponents of these methods,

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including ourselves, hypothesized that under the guidance of mechanical stimuli and the new microenvironment, this tissue capsule will transform into a fully functional artery once implanted into the vasculature (Efendy et al., 2000). The foreign body material (called mandril) serves as a mold, which is removed before vascular grafting. Such a method takes advantage of the plasticity of tissues since it appears that it is the cell and matrix environment rather than their source that determines eventual function (Edelman, 1999). In addition, major limitations of in vitro tissue engineering approaches are circumvented. Indeed, no time-consuming cell culture steps are required. Furthermore, a major difficulty with any in vitro vascular engineering approach is that cells in culture alter their phenotype as well as their immunogenic properties (Thomas and Campbell, 2001). For example, VSMCs in culture modulate from a contractile into a synthetic phenotype (Chamley-Campbell et al., 1979) whereas healthy arteries are composed of contractile VSMCs. Sparks’ mandril Many years ago, Sparks (1973) published a revolutionary in vivo method for in vivo tissue engineering of vascular grafts. His method consisted of preparing a silicone rod of desired diameter and length with coverings of Dacron mesh tubes. This mandril was implanted at the location of the contemplated arterial grafting procedure. After a waiting period of 5 to 8 weeks, during which an autologous tube grew around the silicone mandril and encapsulated the knitted Dacron tubes within its wall, the mandril was withdrawn and the new Dacron reinforced autologous tube was implanted as arterial bypass graft. In vivo studies in dogs showed promising results but clinical studies using these TEBVs as femoropopliteal bypass grafts failed, mainly due to unacceptable high rates of thrombosis and aneurysm formation (Roberts and Hopkinson, 1977). Most likely, the latter failure relates to the absence of an endothelial, anti-thrombogenic layer at the luminal site of the graft and the lack of elasticity (Opitz et al., 2004). Peritoneal cavity as an in vivo bioreactor Ten years ago, we proposed the concept of using the peritoneal cavity as a bioreactor in which to grow an ‘artery’ (Campbell et al., 1999). This site not only serves as a humidified bioreactor in vivo, but also supplies the source of cells, nutrients, growth factors, and other necessary components to create new arteries. Initial studies in rats and rabbits showed that the implanted foreign bodies in the peritoneal cavity became encapsulated with cells of hemopoietic origin (Campbell et al., 2000). Subsequent studies in dogs revealed optimal tissue capsule formation when polyethylene

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(c)

L

5.2 Morphology of TEBVs obtained from the peritoneal cavity. (a) Central region of a TEBV, 6.5 months after it had been grafted into the femoral artery in a dog. Photomicrograph of immunostaining for (b) smooth muscle myosin and (c) Ulex lectin of a TEBV, 4 months after grafting into the femoral artery of a dog. In (b) and (c), the reactive cells (which express the antigen) are stained white. Smooth muscle myosin and Ulex lectin are markers for VSMCs and ECs, respectively. In (b) and (c), the luminal side of the tissue is depicted with the letter L.

tubing encapsulated by Dexon mesh were inserted (Chue et al., 2004). Within 2–3 weeks after insertion of the tubing, the encapsulating cells gradually trans-differentiated into myofibroblasts that subsequently synthesized a collagen matrix, forming a capsule of living tissue covered by a single layer of mesothelium on the outer surface. Using a transgenic mouse model, a PhD student in our laboratory, Jane Mooney, has recently shown the origin of these cells to be monocyte/macrophages (Mooney et al., 2010). After removal of the polyethylene tubing, the tissue capsules were transplanted as interposition grafts in the femoral artery of the same animal in which they were grown. By doing so, the capsules transformed to arteries that consist of ECs, VSMCs, fibroblasts and extracellular matrix proteins, including collagen and elastin (Fig. 5.2). At time of harvest, between 3 and 6.5 months, 73% of TEBVs remained patent. The presence of elastin in this engineered vessel after grafting is particularly encouraging, because the lack of elastin in engineered grafts is believed to cause late dilatation of engineered conduits in high-pressure arterial circuits (Opitz et al., 2004) and thus may contribute to aneurysm formation. Similar results were obtained when the TEBVs were grafted in the aorta of rats (Campbell et al., 1999) or in the carotid artery of rabbits (Campbell et al., 2000). Functionally, these grafts revealed remarkable biocompatibility, together with burst strengths of >2500 mmHg, elasticity and vascular reactivity (Chue et al., 2004). By 3–4 months, the artificial arteries in the high-pressure sites have doubled in thickness, and an ‘adventitia’ containing vasa vasorum has developed on their outer surface (Thomas et al., 2003). Currently, we are evaluating the efficacy of these artificial arteries in improving patency of arteriovenous grafts in pigs. These studies will serve

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as the litmus test before human trials can be initiated. Subsequently, it remains to be tested whether similar tubes of living tissue can be grown in the peritoneal cavity of humans, particularly those with conditions such as diabetes mellitus or chronic renal failure. Indeed, recent studies showed impaired progenitor cell function in patients with these diseases (Loomans et al., 2004; Westerweel et al., 2007).

5.3.6 Use of stem cells for vascular tissue engineering The rapid expansion of knowledge about stem cell biology provided an enormous boost for new studies in the field of (vascular) tissue engineering. Stem or progenitor cells are characterized by their capacity to differentiate into specialized cell types and they can be found in the embryo or adult. Human embryonic stem cells have ethical constraints and the desire for autologous grafts make adult progenitor cells more suitable for tissue engineering. Adult stem cells have been isolated from a variety of sources, such as blood, bone marrow and fat (Sales et al., 2005). In view of their potential to differentiate into vascular cells, endothelial progenitor cells (EPCs) (Asahara et al., 1997) and smooth muscle progenitor cells (SPCs) (Saiura et al., 2001) have been utilized for seeding of prosthetic grafts (see Section 5.3 ‘cell seeding of prosthetic grafts’) and xenogeneic decellularized scaffolds. Kaushal et al. (2001) showed encouraging results of decellularized porcine vascular scaffolds that were seeded with circulating EPCs, and subsequently implanted as a xenogeneic interposition graft in sheep. While all (n = 4) non-seeded conduits thrombosed, all (n = 7) EPC-seeded were patent after 130 days of follow-up. Despite their successful studies in sheep, no reports have been published on EPC-seeded xenogeneic scaffolds in humans thus far. Cho et al. (2005) harvested canine bone marrow-derived progenitor cells which were differentiated into VSMCs and ECs in vitro. Subsequently, these cells were seeded onto decellularized canine carotid arteries. Eight weeks after implanting these TEBVs as interposition grafts in carotid arteries in dogs, all grafts remained patent and showed remarkable regeneration of the three-layered structure of native arteries. Although these animal studies are very encouraging, the authors have not yet attempted to replicate this success in human studies. In addition, none of these studies has shown that stem cells may enhance elastogenesis, which is a prerequisite for successful application of TEBV in the systemic arterial circulation. In contrast, stem-cell seeded scaffolds have already been implanted into the pulmonary circulation in humans. As already mentioned, the low blood pressure in the pulmonary circulation may favor successful application of TEBVs since the biomechanical properties of TEBV are not as high as are required for the systemic arterial circulation. In 2001, researchers at The

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Heart Institute of Japan started to use bone-marrow-cell-seeded PCLbased scaffolds for pulmonary artery reconstruction in pediatric patients (Matsumura et al., 2003). The authors reported on successful implantation of TEBVs in 17 patients. No complications such as thrombosis or stenosis were observed. Long-term follow-up studies should reveal the ultimate clinical value of this technique. For future studies, mesenchymal stem cells (MSCs) could serve as an alternative source for vascular tissue engineering since differentiation of these cells into VSMCs (Narita et al., 2008) and ECs (Yue et al., 2008) has recently been demonstrated.

5.4

Future trends

5.4.1 Validation in relevant large animal models In medical research, there is a tendency to suggest a potential breakthrough once short-term results of new therapies are obtained. The same is true for tissue engineering approaches. However, long-term follow-up studies in relevant large animal models are a prerequisite before clinical trials are initiated. For cardiovascular studies, pigs are frequently used as models because of their analogous vascular anatomy, size, and physiology (Ferrell et al., 1992). In our opinion, the arteriovenous graft model is the optimal application for preclinical and clinical evaluation because (i) there is an urgent need for better vascular accesses, (ii) graft failure is unlikely to be life threatening and (iii) it is the most challenging model with respect to the mechanical environment. The latter relates to the turbulent flow pattern and blood flow rates usually exceeding 800 mL/min, the gradual decline in blood pressure along the course of the graft and the repetitive punctioning of the graft during hemodialysis sessions (Rotmans et al., 2005b). With regard to these animal studies, it is important to notice that elasticity of arteries does vary between species. Porcine coronary and internal mammary artery walls were found to be three times more elastic than human arteries (van Andel et al., 2003). The latter may result in underestimation of wall stress and the risk of rupture when TEBVs are evaluated in pigs.

5.4.2 Hurdles on the way to widespread clinical use Despite the significant improvements that have been achieved in the last two decades, the ideal substitute for small-diameter arteries still remains to be constructed. For future studies, elastic compliance is one of the most crucial characteristics of TEBVs that has to be ameliorated. Not surprisingly, most TEBVs that were generated under static conditions lack tensile strength and elasticity. Indeed, several studies showed that shear stress is an important stimulator of collagen (Jin et al., 2001) and elastin (Isenberg

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and Tranquillo, 2003) synthesis and induces the circumferential orientation of VSMCs (Seliktar et al., 2000). In subsequent efforts to design an ideal tissue engineered blood vessel, many investigators examined different types of mechanical stimulation techniques using bioreactors in order to mimic the exposure to shear stress in vitro. Although most of them succeeded in improving burst strength as a result of collagen synthesis, most failed to induce elastin synthesis (Niklason et al., 1999; Hoerstrup et al., 2002). Biosynthesis and subsequent crosslinking of elastin appears to be one of the most complex and tightly regulated processes during the maturation of blood vessels (Bunda et al., 2005). Incorporation of this knowledge about elastin metabolism in new strategies for vascular tissue engineering is a prerequisite to improve tissue elasticity. Another important issue that has to be addressed in future studies is the formation of vaso vasorum in TEBVs to ensure sufficient nutritional supply of the thickening vascular construct. Indeed, if oxygen concentrations are inadequate, cell proliferation ceases and cell viability begins to break down (Kellner et al., 2002). In the past, the adventitial segment of the vessel wall has received limited attention compared with the endothelium. However, the adventitia (in which the vaso vasorum is located) has emerged as an active participant in vascular homeostasis and remodeling (Sartore et al., 2001). Therefore, adventitial delivery of progenitor cells, vascular growth factors or gene therapeutic constructs may enhance the functional capacity of this important layer of TEBVs. Besides the above-mentioned anatomical and functional requirements, a limited time frame required for the production of TEBVs is essential for future large-scale clinical application. Therefore, time consuming in vitro cell culture procedures need to be reduced as much as possible. In this respect, it remains to be determined if EC seeding of engineered constructs prior to implantation is a prerequisite for successful application. Although prior cell seeding has obvious advantages for immediate graft function, cellular retention after implantation is usually low (Bhat et al., 1998). Omitting this step could therefore save a substantial amount of valuable time. Evidence for potential success of such an approach comes from human (Gravanis and Roubin, 1989) and animal studies (Reidy and Schwartz, 1981) that revealed spontaneous reendothelialization of denudated arterial segments, for instance after balloon angioplasty. Furthermore, some decellularized xenogeneic scaffolds (Bergmeister et al., 2008) have shown to endothelialize spontaneously after implantation into the arterial circulation. Additional modifications that increase mobilization and homing of EPCs may further enhance the formation of an endothelial layer on the luminal surface of TEBVs in vivo. For instance, incorporation of growth factors such as transforming growth factor-β and vascular endothelial growth factor into engineered constructs may improve endothelialization,

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since these growth factors have emerged as pivotal regulators of EPC function (Hristov et al., 2003). Such efforts to facilitate in vivo endothelialization may kill two birds with one stone since laborious in vitro seeding procedures are circumvented while the full regenerative potential of vascular progenitor cells can be exploited. As a result of recent advances of the field of regenerative medicine, it has been hypothesized that the ultimate configuration of an autologous cell-based vascular graft need not be determined at outset by the cells that comprise the device, but rather by the dynamic environment, wherein the body modifies the tissue-engineered construct to meet local flow, metabolic and inflammatory requirements (Edelman, 1999).

5.5

Conclusion

Vascular tissue engineering is a rapidly expanding field and represents the newest concept in order to alleviate the limitations of small-diameter prosthetic grafts. Although the use of human TEBVs in the pulmonary circulation has shown inspiring clinical success in pediatric patients, the promise of an ‘off-the-shelf’ tissue-engineering graft for adult revascularization remains unrealized. When compared to other organs, such as the kidney, the structure of blood vessels appears relatively simple. However, despite more than twenty years of research, the ‘Holy Grail’ is still to be discovered. One hundred and fifty years after the publication of ‘On the Origin of Species’ by Charles Darwin, that work should not only result in enormous admiration for the evolution of mankind, but should also encourage the scientific community to continue its efforts to create autologous vascular substitutes. In this respect, the broadly supported focus of stem cell biology, regenerative medicine and tissue engineering in medical research is encouraging. In the past, progress in tissue engineering may have been stunted by the existence of separate worlds in which chemical engineers and biomedical scientists have been working. To reach the ultimate goal of manufacturing adequate tissue engineered blood vessels within the foreseeable future, strong collaboration of engineers, biologists and clinicians is of vital importance.

5.6

References

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6 Stem cells for organ regeneration K. D. D E B, Dayananda Sagar Institutions, India

Abstract: Tissue engineering aims at regenerating and restoring organ function by exploiting the self-healing properties of body tissues. To achieve this, cells from different sources, growth inducing molecules, and a biomaterial-support called a scaffold are used in isolation or in combination to reconstruct organs. However, regeneration of a functional tissue or organ has so far remained a challenge. With the recent advances in adult- and embryonic-stem cells technologies, the dream of reconstructing an entire organ or at least parts of it, seems likely to be realized soon. Moreover stem cells provide tissue engineers the freedom to choose cells at a more defined or intermediate stage of differentiation. Such advances will lead to better integration of the tissue engineered grafts in the host. Organ re-construction is performed either in vitro in laboratories and may be made available on demand; or the regenerating factors or microenvironment can be engineered in vivo for organs to grow inside the body. Integration of stem cells and tissue engineering has resulted in tremendous progress, poised to bring a paradigm shift in regenerative medicine. This chapter presents an overview of the basics and the major advances made in this area so far. Key words: stem cell, organ regeneration, tissue engineering, scaffolds, biomaterials.

6.1

Introduction

Advances in modern medicine have evolved from multidisciplinary research across a wide range of scientific and engineering disciplines. Tissue engineering is defined as an ‘interdisciplinary field that applies the principles of engineering and life sciences toward the development of biological substitutes that restore, maintain or improve tissue function’ (Langer and Vacanti, 1993; Nerem, 1991). It is based on principles to exploit the self-healing potential of an organ to regenerate itself. Increase in average life span of human beings over the last couple of decades has resulted in a vast proportion of patients suffering from agerelated degenerative diseases, organ failures, congenital malformation or other debilitating conditions. Until recently, treatment options for such patients were limited to either prosthetic implants or organ transplantations. In the first approach, reconstructive surgery is performed by inserting 147 © Woodhead Publishing Limited, 2010

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an artificial organ substitute, also called a biomedical device. Examples include artificial pacemakers, heart valves, gold teeth, fracture fixation plates, kidney hemodialyzers and prosthetic joints, to substitute for the damaged organs. However, in most of these cases the artificial organs cannot perform all the functions of the organ, and lead to progressive deterioration of the health of the patient (Tabata, 2000). Moreover, many of these devices lack lifelong durability. In an alternative approach, an organ obtained from a live donor or cadaver is transplanted into the patient. Though mostly successful, such transplantations are limited by the number of matching donors available. Both implanted devices and transplanted organs have to face immune rejection from the patient. The advent of immunosuppressive drugs such as glucocorticoids, cyclosporine, and later azathioprine and monoclonal antibodies, did revolutionize the practice of transplantation medicine to a certain extent (Italia et al., 2006), and enabled organ substitutes to overcome foreign body rejections. These drugs, however, have severe adverse effects for long-term therapy and predispose patients to recurrent infections. Hence, the clinical solutions for organ failure so far remain imperfect. Organogenesis is induced to occur in laboratories and transplanted into the individuals, or the necessary factors are delivered in the injury site in vivo. This is achieved by engineering in tandem one or more of the three elements of tissue engineering – cells, tissue-inducing factors and materials – to support and guide tissue growth (Langer and Vacanti, 1993). It can be implied that a typical tissue engineer has to work at the interface of biological, chemical, physical, engineering, medicine, genetics, pharmaceutical, and materials sciences to achieve the desired goals. Tissue engineering is therefore a relatively modern branch of science, as compared to the allied practices of regenerative medicine and cell-based therapies. Regeneration and stem cells have been widely studied subjects, particularly in insect and other organisms. The ability of lizards to regenerate lost tails is appreciated with amazement by most of us during our childhood. According to the Book of Genesis, 2:21, a rib was harvested from Adam, the first donor, and was used to fashion Eve. In Greek mythology a story goes that Prometheus was to be punished for giving fire to mankind, and an eagle was sent to devour his liver; however, he survived because he could quickly regenerate his liver every day (Rosenthal, 2003). In Ancient India, Ayurvedic texts dating back to 600 bc contain a detailed description of protocols to carry out rhinoplasty, a procedure intended to resurrect an injured nose with skin obtained from other parts of the body. Shusruta, the surgeon who described such a procedure, is regarded as a pioneer of skin grafting and plastic surgery in India (Rana and Arora, 2002). In modern times, routine graftings with cornea were performed in Vienna during the

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early years of the 19th century (Gardner, 2007). In the 1930s, a Swiss physician, Dr Paul Niehans, also known as the father of cell therapy, injected a suspension of steer thyroid cells into a critical patient suffering from thyroid deficiency. The patient went on to live for the next 30 years and this heralded the world of modern cell therapy (www.healthatoz.com/celltherapy). The clinical success of the first ‘artificial kidney dialyzers’ in 1944 (Kolff et al., 1944) ushered in the era of engineering sciences’ entry into healthcare research. The next big breakthrough in regenerative medicine was the first successful heart transplantation by Dr Christiaan Barnard in the mid-1960s (Barnard, 1967). Thereafter, in the 1970s artificial pacemakers were successfully introduced in clinical applications (Niklason and Langer, 2001). This was followed by approval of cyclosporine for clinical use as an immunosuppressant in 1983. These events together brought great strides in the management of patients affected with some forms of organ failure. However, concerns over the long-term use of artificial organs and immunosuppressive drugs surfaced soon, and unavailability of matching donors made the situation no better. In modern times, the earliest applications of tissue engineering in clinical practice started when Drs Bruke, Yannas and Bell prepared collagen-based matrices, seeded with dermal fibroblasts, to treat burn victims, in the early 1980s. The first in vitro skin grafting was successfully tested clinically at around the same time (Yannas et al., 1982). Since then, tissue engineering has made a formal entry in the scientific literature, and became more popular during the late 1980s and early 1990s.

6.2

Basic components of tissue engineering

Tissue engineering for organ regeneration mainly utilizes various isolated cell types and aims at creating a structurally and functionally active organ system that can substitute for the damaged tissue in vivo. As mentioned earlier, tissue engineers try to use cells, tissue regenerating factors and scaffolding materials to create new tissues. These are the three primary elements of tissue engineering. A tissue engineer has to select a right mix and a formula comprising these components and apply them in conjunction to fulfill the objectives. With the advent of stem cell technologies and isolation procedures, tissue engineers are now trying to harness the ability of stem cells to self-renew and differentiate to various lineages.

6.2.1 Cell sources and types Cells available for tissue engineering are classified in various ways. Depending upon the donor of the cells, there are three types of tissue engineering – autologous, allologous and xenologous. In Dr Niehans’ experiment

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described in Section 6.1, he obtained cells from a different species, the steer, and used them to restore a patient’s dysfunctional organ. The procedure of introducing cells obtained from an individual of a different species is called xeno-grafting. Xeno-grafts are susceptible to immune reactions and may produce the challenge of cross infections (Appel et al., 2000; Platt, 2000). However, as some studies have shown, cross infectivity is always a possibility and not a necessity (Levy et al., 2000). Apart from this, the host will identify the xeno-graft as a foreign body and will try to reject these cells, resulting in inflammation and so called graft-versus-host reactions. Thus, xeno-grafting is accompanied by supplements of immunosuppressive drugs such as cyclosporine. If tissue engineering is performed utilizing cells of an individual of the same species as the patient, it is called allologous grafting. It is desired that the donor and the recipient share the same human leukocyte antigen for graft survival and cell acceptance. However, it is often very difficult to obtain matching donors, and immuno-modulators may be required in these procedures. This brings us to the most desirable situation for tissue engineering, at least from an immunological point of view. In this instance, cells harvested from a patient himself are induced to grow and substitute the distorted organ. This can be accomplished either in vitro or in vivo. Such a procedure is known as auto-grafting or autologous tissue engineering. An alternative classification of cells available for tissue engineering is based upon their morphological and physiological characteristics. Thus, tissue engineers have the choice of using: (i) tissue-specific differentiated cells, (ii) adult progenitor or adult stem cells, (iii) embryonic stem cells, and the more recently created (iv) induced pluripotent (iPS) stem cells. The following sections cover each of these cell types. Somatic cells In vitro expansion of cells to significant quantities has been a limitation when developing cell-based regenerative medicine techniques for organ replacement. Though some organs, such as the liver, have a high regenerative capacity, hepatocyte growth and expansion in vitro can be difficult. By studying the privileged sites for committed precursor cells in specific organs, as well as exploring the conditions that promote differentiation and/or selfrenewal of these cells, it may be possible to overcome the obstacles that limit cell expansion in vitro. Organ regeneration and tissue engineering attempts with most adult cell types have met with varying degrees of success. Adult autologous tissue-specific cells are harvested from the patient, expanded ex vivo and reintroduced into the body. A limitation of using mature cells for tissue engineering is, with the exception of some cell types such as the dermal fibroblasts, keratinocytes, and some epithelium cells such

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as corneal epithelia, cartilage and adipocytes, harvesting adult tissuespecific cells is difficult and requires invasive surgeries, if not biopsies. Such surgeries may result in donor site cell morbidity and scarring. The two most developed autologous mature cell therapies that have advanced from the laboratory to the clinic involve the repair of cartilage using autologous chondrocytes, and the treatment of burns with autologous cultured keratinocytes (Fodor, 2003). A few tissue engineered products based on allologous adult mature cells such as Apligraft (composed of neonatal foreskin keratinocytes and dermal fibroblasts), are available in the market and many others are actively being investigated (Chu et al., 1995; Eaglstein and Falanga, 1998; Kondziolka et al., 2000). Xenologous adult cells are likely to pose significant challenges. Clinical trials with transgenically engineered pig livers to detoxify the blood of patients suffering from fulminant hepatic failure (FHF) via extracoporial perfusion has been conducted (Levy et al., 2000). Tissue engineering with allologous and xenologous cells need to apply technologies such as cell encapsulation for immune isolation, modified tissue processing methods, tolerance induction in patients and production of transgenic organisms with humanized antigenic properties to overcome their inherent disadvantages. Galα1,3 gal transferase null transgenic pigs have already been produced and they represent a significant development towards eliminating both hyperacute and acute vascular rejection and extended survival of xenologous cells and organs (Phelps et al., 2003). Despite the advent of stem cells, adult cells to some extent continue to excite tissue engineers in their pursuit to create new tissues (Kucia et al., 2007). Somatic stem cells Adult stem cells have been known for more than two decades now; however, it is only the recent advances in our understanding of stem cells that offer the promise to convert all the expectations of tissue engineering into reality, and revolutionize the practice of regenerative medicine. The reasons why stem cells appear to be such ideal candidates for regenerative medicine and tissue engineering are attributed to their characteristic ability to self-renew and undergo multiphenotypic differentiation. There are three types of stem cells – embryonic, germinal and somatic/adult/foetal stem cells. The potential of stem cells to give rise to a number of different cell types determines their therapeutic utility. Thus stem cells are classified as pluripotent (embryonic stem cells or ES cells), multipotent (adult stem cells), and unipotent (epidermal stem cells) (Serakinci and Keith, 2006). Thus, discounting other factors, ES, which can get differentiated to a wider number of cell phenotypes compared to adult stem cells, should be better candidates for tissue engineering. However, adult stem cells (except neural

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stem cells) are easier to harvest. Almost all organs of the adult human body contain a specific mass of cells possessing the capacity to self-renew and differentiate. These cells may be quiescent in some tissue beds while active in others. The bone marrow is one of the most abundant sources of selfreplicating mesenchymal and hematopoetic stem cells, both of which are multipotent in character. Human MSC can give rise to a variety of tissues such as adipocytes, bone, cartilage and muscles, and can be used for tissue engineering purposes. In addition, neuronal stem cells isolated from the brain or the spinal cord are relatively easy to expand in culture and can be used for various clinical ailments such as Parkinson’s, Alzheimer’s and other brain trauma. However, adult stem cells lack the immortality of ESCs – a desirable feature for tissue engineering. This feature can be introduced by genetic transductions of cells, e.g. ectopic expression of the human telomerase reverse transciptase gene may render them capable of infinite proliferation (Natesan, 2005). Recently it has been shown that, under certain conditions, mature adult cells can be induced to form a more primitive phenotype by a process termed ‘dedifferentiation’. This technology is discussed in more detail in the section for induced pluripotent cells (iPS), which follows later. Adult stem cells isolated from a particular tissue can also form cells of another phenotype, this phenomenon being known as ‘trans-differentiation’. For example, co-culturing neural stem cells with muscle progenitor cells has been shown to form muscle without the need for any other factor (Galli et al., 2000). Both dedifferentiation and trans-differentiation, when established completely, will greatly increase the clinical applicability of adult stem cell therapy (Oliveri, 2007; Thowfeequ et al., 2007). Adult stem cells derived from adult bone marrow and muscles are currently being used by several workers for tissue engineering experiments (Wei et al., 2007). Besides this, it is also believed that adult tissues harvest a small quiescent population of pluripotent stem cells known as the embryonic-like stem cells (ELSCs) (Young and Black, 2004). It is thought that, during development, a small number of cells escape the development continuum, migrate and home-in at various parts/tissues of the adult. We have seen that the umbilical cord Wharton’s jelly is a rich source of such population of stem cells (unpublished data). Also, the placental and amniotic fluid stem cells have been shown to be very rich sources of adult multipotent stem cells. The following section presents a brief description of these cells. Placental and amniotic fluid stem cells Both amniotic fluid and placenta are known to contain multiple partially differentiated cell types derived from the developing fetus. These cells can be obtained from amniocentesis, or from chorionic villous sampling in the

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developing fetus, or from the placenta at the time of birth. Stem cell populations from these sources, called amniotic fluid and placental stem cells (AFPSCs), that express embryonic and adult stem cell markers have been successfully isolated (DeCoppi et al., 2007). The undifferentiated stem cells expand extensively without the need for a feeder cell layer and double every 36 hours. Unlike human embryonic stem cells, the AFPSCs do not form tumors in vivo. Lines maintained for over 250 population doublings retained long telomeres and a normal karyotype. These cells can be differentiated into neuronal lineage secreting the neurotransmitter Lglutamate or expressing G-protein-gated inwardly rectifying potassium (GIRK) channels, hepatic lineage cells producing urea, and osteogenic lineage cells forming tissue engineered bone. Embryonic stem (ES) cells and derivatives In 1981, pluripotent cells were found in the inner cell mass of the human embryo, and the term ‘human embryonic stem cell’ was coined (Martin, 1981). These cells are known to differentiate into all cells of the human body, excluding placental cells (only cells derived from the morula are totipotent; that is, able to develop into all cells of the human body). These cells have great therapeutic potential, but their use is currently limited by several factors, both biological and ethical. The political controversy surrounding ES cells started in 1998 with the creation of human ES cells from discarded, non-transferred human embryos. Human ES cells were isolated from the inner cell mass of a blastocyst (an embryo five days postfertilization) using an immunosurgical technique where complement proteins and antibodies lyse the trophectoderm so that only the inner cell mass survives. Given that some cells cannot be expanded ex vivo, ES cells could be the ideal resource for tissue engineering because of their fundamental properties: the ability to self-renew indefinitely and the ability to differentiate into cells from all three embryonic germ layers (Deb et al., 2008). These cells have demonstrated longevity in culture and can maintain their undifferentiated state for at least 80 passages when grown using current published protocols. During the past few years, ES cells have evolved as a strong contender for cell therapies and tissue engineering (Deb and Sarda, 2008). These cells offer unique advantages over other cell types for therapy, and some of these therapies based on ES cell progenitors are now close to approval. While protocols for derivation of almost every type of cell from mouse ES cells have been published, derivation of many different cell phenotypes from human ES cells are also being worked out. Human ES cells can give rise to neurons, cardiomyocytes, pancreatic islet cells, hematopoeitic cells, RBCs, etc., under controlled culture conditions. Guided differentiation of human

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ES cells towards dopaminergic and motor neurons in the laboratory is made possible (Wichterle et al., 2002). ES cells allow tissue engineers to define a specific stage of differentiation for transplantation. ES cells also exhibit features that bestow upon them unlimited potential for self-renewal, thus making them prime candidates for cell banking (Mitjavila-Garcia et al., 2005). Banked cells can then be used for therapy when needed. A focus of current research on ES cells is to provide karyotypic stability over repeated passages (Buzzard et al., 2004). Human ES cells cultured on mouse fibroblast feeder cells pose a threat for xenogenic contamination and are therefore unsafe for clinical applications. Through modern tissue engineering, considerable efforts are being directed towards eliminating feeder layers so as to make them more useful for therapeutic purposes (Amit et al., 2004). However, clinical application of ES cells is limited because they represent an allogenic resource and thus have the potential to evoke an immune response. Beside this, the ES cells, when injected, give rise to teratocarcinomas, and this has so far remained a major hurdle. Better differentiation protocols with higher efficiency are required to obtain terminally differentiated derivatives before application in regenerative medicine. Also, there is a need to develop technologies to sort out undifferentiated populations of ES cells following a terminal differentiation protocol (Deb and Sarda, 2008). Alternative and new stem cell technologies (such as somatic cell nuclear transfer and reprogramming) hold promise to overcome the current limitations.

6.2.2 Therapeutic cloning: somatic cell nuclear transfer Somatic cell nuclear transfer (SCNT) is a technique used to derive patientspecific pluripotent stem cells. SCNT basically means the removal of an oocyte nucleus in vitro, followed by its replacement with a nucleus derived from a somatic cell obtained from a patient. Activation of this oocyte to stimulate cell divisions up to the blastocyst stage, is then attained either with chemicals or electricity. At this time, the inner cell mass is isolated and cultured, resulting in ES cells that are genetically identical to the patient. It has been shown that nuclear transferred ES cells derived from fibroblasts, lymphocytes, and olfactory neurons are pluripotent and generate live pups after tetraploid blastocyst complementation, showing the same developmental potential as fertilized blastocysts (Byrne et al., 2007; Meissner et al., 2007; Mitalipov, 2007; Wernig et al., 2007). Thus, the resulting ES cells are a perfect match to the patient’s immune system and would prevent rejection. Although SCNT-derived ES cells contain the nuclear genome of the donor cells, the mitochondrial DNA (mtDNA) contained in the oocyte could lead to immunogenicity after transplantation. Lanza et al. carried out an elegant experiment to assess the histocompatibilty of nuclear

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transfer-generated tissues. They microinjected the nucleus of a bovine skin fibroblast into an enucleated oocyte, and blastocysts were generated (Lanza et al., 2002). The blastocysts were implanted (reproductive cloning), with a purpose of harvesting renal, cardiac and skeletal muscle cells, which were then expanded in vitro, and seeded onto biodegradable scaffolds. These scaffolds were then implanted into the donor from whom the cells were cloned, to determine if the cells were histocompatible. The experiment revealed that cloned renal cells showed no evidence of T-cell response, suggesting that rejection would not necessarily occur in the presence of oocytederived mtDNA. These findings demonstrate the power of SCNT in overcoming the histocompatibility problem of stem cell therapy. Although promising, SCNT has certain technical and ethical limitations that require further improvement before its clinical application. The ethical concern is pertaining to the potential of the embryos resulting from SCNT to develop into cloned embryos if implanted into a uterus. Many animal studies have shown that blastocysts generated from SCNT can give rise to a liveborn infant that is a clone of the donor. In 1997, for example, a sheep named Dolly was derived from an adult somatic cell using nuclear transfer. This process is known as reproductive cloning, which is banned in most countries for human applications. In contrast, therapeutic cloning is used to generate only ES cell lines whose genetic material is identical to that of their source, and the generated blastocysts are never implanted into a uterus. In this case, blastocysts are only allowed to grow in culture until they reach a 100-cell stage, from which ES cells can be obtained. Recently, nonhuman primate ES cell lines were generated by SCNT of nuclei from adult skin fibroblasts (Byrne et al., 2007; Mitalipov, 2007). A total of 304 oocytes yielded 35 blastocysts, from which two ES cell lines were derived. Both lines demonstrated typical ES cell morphology. They also demonstrated selfrenewal and expressed the stem cell markers OCT4, SSEA4, LEFTYA, TDGF, TRA1-60 and TRA1-80. To test their differentiation potential, the cells were exposed to cardiac and neural differentiation conditions, and these experiments resulted in cells that expressed markers of the specified lineages. When injected into SCID (severe combined immunodeficiency) mice, the SCNT-derived non-human primate ES cells induced teratomas that contained differentiated cell types from all three embryonic germ layers. A careful assessment of the quality of the lines must be done before SCNT-derived ES cells can be used for therapy. Some cell lines generated by SCNT have been shown to contain chromosomal translocations and it is not known whether these abnormalities originated from aneuploid embryos or if they occurred during ES cell isolation and culture. In addition, the low efficiency of SNCT (0.7%) and the inadequate supply of human oocytes further limits the therapeutic applications of this technique.

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6.2.3 Reprogrammed somatic induced pluripotent stem (iPS) cells Very recently, many exciting reports of the successful transformation of adult cells into pluripotent stem cells through various kinds of genetic ‘reprogramming’ have been published. Reprogramming is a technique used to de-differentiate adult somatic cells to produce patient-specific pluripotent stem cells. Yamanaka was the first to discover that mouse embryonic fibroblasts (MEFs) and adult mouse fibroblasts could be genetically modified and reprogrammed into an ‘induced pluripotent state’ (iPS) (Takahashi and Yamanaka, 2006). Cells generated by reprogramming would be genetically identical to the somatic cells (and thus, the patient who donated these cells) and therefore would not be rejected. Yamanaka’s group used MEFs engineered to express a neomycin resistance gene from the Fbx15 locus, a gene specifically expressed only in ES cells. They examined twenty-four genes that were thought to be important for embryonic stem cells and identified four key genes that, when introduced into the reporter fibroblasts, resulted in drug-resistant cells. These were Oct3/4, Sox2, c-Myc, and Klf4. This experiment indicated that expression of the four genes in these transgenic MEFs led to expression of a gene specific for ES cells. The resultant iPS cells possessed the immortal growth characteristics of self-renewing ES cells, expressed genes specific for ES cells, and generated embryoid bodies in vitro and teratomas in vivo. When iPS cells were injected into mouse blastocysts, they contributed to a variety of cell types. However, although iPS cells selected in this way were pluripotent, they were not identical to ES cells. Unlike ES cells, chimeras made from iPS cells did not result in full-term pregnancies. Gene expression profiles of the iPS cells showed that they possessed a distinct gene expression signature that was different from that of ES cells. In addition, the epigenetic state of the iPS cells was somewhere between that found in somatic cells and that found in ES cells, suggesting that the reprogramming was incomplete. These results were improved significantly by Wernig et al. (2007). Fibroblasts were infected with retroviral vectors and selected for the activation of endogenous Oct4 or Nanog genes. Results from this study showed that DNA methylation, gene expression profiles, and the chromatin state of the reprogrammed cells were similar to those of ES cells. Teratomas induced by these cells contained differentiated cell types representing all three embryonic germ layers. Most importantly, the reprogrammed cells from this experiment were able to form viable chimeras and contribute to the germ line like ES cells, suggesting that these iPS cells were completely reprogrammed. Wernig et al. observed that the number of reprogrammed colonies increased when drug selection was initiated later (day 20 rather than day 3 post-transduction). This suggests that reprogramming is a slow and

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gradual process and may explain why previous attempts resulted in incomplete reprogramming. It was recently shown that reprogramming of human cells is possible (Takahashi et al., 2007; Yu et al., 2007). Yamanaka’s group showed that retrovirus-mediated transfection of Oct3/4, Sox2, Klf4, and c-Myc generates human iPS cells that are similar to hES cells in terms of morphology, proliferation, gene expression, surface markers, and teratoma formation. Thompson’s group showed that retroviral transduction of OCT4, SOX2, NANOG, and LIN28 could generate pluripotent stem cells without introducing any oncogenes (c-Myc). Both studies showed that human iPS were similar but not identical to hES cells (Takahashi et al., 2007; Yu et al., 2007). Another concern is that these iPS cells contain three to six retroviral integrations (one for each factor), which may increase the risk of tumorigenesis. It was found that tumor formation occurred in chimeric mice generated from Nanog-iPS cells and also that 20% of the offspring developed tumors due to the retroviral expression of c-Myc. An alternative approach using a transient expression method, such as a non-integrating adenovirusmediated system, was successfully used to produce iPS cells recently. Both Jaenisch and Yamanaka showed strong silencing of the viral-controlled transcripts in iPS cells (Okita et al., 2007; Meissner et al., 2007). This indicates that these viral genes are required only for the induction, not the maintenance, of pluripotency. Another concern about the findings was the use of transgenic donor cells for reprogrammed cells in the mouse studies. Both studies used donor cells from transgenic mice harboring a drug resistance gene driven by Fbx15, Oct3/4, or Nanog promoters so that if these ES cell-specific genes were activated, the resulting cells could be easily selected using neomycin. It was important to assess whether iPS cells could be derived from nontransgenic donor cells, for applications in therapy. Wild type MEF and adult skin cells were therefore retrovirally transduced with Oct3/4, Sox2, c-Myc, and Klf4 and ES-like colonies were isolated by morphology alone, without the use of drug selection for Oct4 or Nanog (Meissner et al., 2007). These iPS cells from wild type donor cells formed teratomas and generated live chimeras. This study suggests that transgenic donor cells are not necessary to generate IPS cells. Although very exciting, there are still many questions about the reprogramming that need to be addressed before these cells become clinically eligible. Use of retroviral and adenoviral constructs for genetically modifying cells may face problems in getting approvals from the regulatory agencies in many countries such as India. Trials are therefore being carried out in our laboratories to reprogram cells by using chemical entities or natural compounds. We are also trying to reprogram cells by changing their niche, and the microenvironment. Micro-encapsulation of the cells using different

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biomaterials is already being carried out in our laboratories to achieve such de-differentiation from somatic cells. Tissue engineering may therefore have a major role to play in this area.

6.2.4 Growth factors For successful application of cells in tissue engineering, it is essential to understand and modulate the key signaling mechanisms involved in cellular differentiation and proliferation. The processes of differentiation and proliferation are controlled by both intrinsic regulators and the extracellular environment. The intrinsic environment consists of a cocktail of growth factors, signaling molecules, and hormones. The signaling pathways are triggered when an appropriate receptor is activated by a ligand. This ligand can be a protein, mechanical stimulus, molecules like retinoic acid (Wichterle et al., 2002) or a synthetic small molecule. Mechanisms of action of most drug molecules in clinical practice today are based on such ligand–receptor interaction. FGF, Wnt, Hedgehog, TGF/BMP, and Notch are some of the signaling pathways responsible for mammalian organogenesis. These pathways are also supposed to play a significant role in tissue repair and regeneration (Wu and Sheng, 2006). Kinase proteins on cell surfaces are postulated to be some of the key receptors for regulation of cell differentiation (Rzucidlo et al., 2007; Sato et al., 2004). Interestingly, protein kinases are favored by medicinal chemists also as potential targets for many disorders (Thaimattam et al., 2007). Thus a few investigators, aided by advances in proteomics and bioinformatics tools, have already started designing and screening libraries of synthetic small molecules that selectively bind to these receptors and bring about morphogenesis of cells into a specific lineage (Sheng et al., 2003). The quintessential role of telomerase in regulating cell proliferation and division is well appreciated (Klinger et al., 2006; Peterson and Niklason, 2007). In the years to come, synthetic molecules which can bind selectively to the signaling receptors bcl-2 and telomerase will become an indispensable tool for the tissue engineer in regenerative therapy. Some of the growth factors used for tissue engineering purpose are fibroblasts growth factor, bone morphogenic protein, insulin-like growth factor, and vascular endothelial growth factor, amongst others.

6.2.5 General scaffolds and materials Neither isolated cells nor growth factors can be relied upon to produce tissues of all practical sizes and shapes. Here, tissue engineering enters the domain of material science. Development of biomaterials for tissue engineering is as exciting a field of research as stem cells and molecular biology. The need of biomaterials can be understood looking at the role of extra

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cellular matrix (ECM) in native tissue environment. This matrix, composed of various biopolymers conjugated with glycol-saccharides and proteins, is responsible for providing mechanical strength and physical support to the growing cells, and directs their proliferation, differentiation, and morphogenesis (Zagris, 2001). It allows cells to assume a particular shape and size. During any tissue damage, a significant amount of ECM is also lost. Scaffolds are those structures that are fabricated to replace ECM in vivo or provide an artificial ECM for in vitro regeneration. It is sometimes possible to regenerate tissues by providing just the scaffold at the site of injury. Thus, cells get the necessary support to grow. As the growing cells can secrete their own ECM, an ideal scaffold should degrade with time, without leaving any scar. The degradation products should be non-toxic. In tissues, where the cells have little inherent regeneration potential, the scaffold itself may not be sufficient. In these cases, scaffolds are seeded with replicative cells and suitable growth factors to form the desired tissue. The characteristics of a good scaffold material are high porosity and proper pore size, high surface-area-to-volume ratio – this allows maximum cell–cell interaction and cell–scaffold interaction, biodegradability, high mechanical strength, and high degree of biocompatibility (Ma, 2004). Below we present some of the materials used in tissue engineering. Metals The impact of metallic implants prevails in orthopedics, dentistry, craniofacial surgeries, etc. Stainless steel, cobalt and titanium based alloys are the metals mostly used in medicine (Brunski, 1996). Inertness coupled with mechanical integrity has made metals suitable for implantation. Though more attention is paid to other types of biomaterials for tissue engineering, metals such as titanium do find continued use as an adjunct material (Liu et al., 2006). Ceramics Ceramics and bio-glasses comprise a range of inorganic materials with established use in dentistry and some medical application such as eye glasses. These materials include alumina, silicon dioxide, TiO2, CaO and P2O5. Some porous ceramics isolated from corals appear to be suitable, for they can effectively provide surface for bone remodeling, especially when load bearing is not the primary criterion (Schors and Holmes, 1993). Amongst all ceramic materials, calcium phosphate has gained widespread acceptability for tissue engineering. Calcium phosphate and hydroxyapetite impregnated with growth factors and antibiotics appear attractive propositions for scaffold fabrication (Nie and Wang, 2007).

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Natural polymers Natural polymers mimic the native ECM down to the molecular level and are better materials for the biological system to recognize. Also, many of them possess the property of self-assembly and hence it is simple to build the ECM in vitro from such polymers. They are degraded into noninflammatory products (a significant advantage over their synthetic counterparts), but these polymers may be immunogenic in nature. Since they are obtained from natural sources, they carry a risk of microbiological contamination, and their exact composition is often difficult to ascertain and reproduce. Many of the components are proteinacious in nature so they are less amenable to the rigors of processing conditions. Collagen, dextran, fibrin, silk, gelatin, sodium alginate, chitosan, and hyaluronic acid are some of such materials that are extensively used for tissue engineering (Yannas, 1993). In fact, collagen was one of the first materials to be applied successfully for clinical tissue engineering (Burke et al., 1981).

Synthetic polymers As a tissue engineer, biomaterial needs to play the role of a temporary scaffold, wherein it is supposed to degrade at a rate corresponding to the kinetics of neo-tissue formation; novel biodegradable materials have gained preference over erstwhile popular synthetic polymers such as polymethylmethacrylate (PMMA) and polyethylene. Linear aliphatic polyesters are amongst the most popular materials for tissue fabrication. Some of their fascinating properties are controlled degradation, and the ability to entrap various drug molecules and yield materials of different strengths. For example, polyglycolic acid (PGA) is a hydrophilic polymer with very fast degradation rate. Polylactic acid, on the other hand, is hydrophobic and persists for over a year in vivo. By blending these polymers in different ratios, materials with customized biological properties can be synthesized. This approach has already evinced much interest in drug delivery, and tissue engineering is fast picking up. Another important advantage is the vast amount of chemical modification that can be applied to these polymers to get desired biocompatibility and cell–material interactions. Polycaprolactone, polyphospharazenes and polyanhydrides are other materials of this class with potential use in tissue engineering (Katti et al., 2002; SokolskyPapkov et al., 2007). Biodegradable polymer scaffolds with interconnected pores allow formation of complex D tissues during differentiation. The scaffold, designed to withstand the compressive strength of cells, provides physical cues for cell orientation and spreading, and promotes differentiation, proliferation and organization of cells. The pores provide space for remodeling of tissue structures (Richardson et al., 2001). Variation of growth

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factor delivery from these scaffolds can lead to formation of desired tissuelike cartilage, and neural and blood vessels. Such engineered tissue-like structures can then be directly used for transplantation (Vacanti and Langer, 1999). From the elements of tissue engineering it can be implied that this multidisciplinary field involves interplay between some of the most exciting fields of modern science, namely, stem cell biology, molecular genetics, protein drug delivery and biomaterials sciences, along with the essential tenets of physics, chemistry, medicine and engineering. Advances in the regeneration of specific tissues and organs using tissue engineering approaches are discussed in the next section.

6.3

Tissue engineering and stem cells in organ regeneration

Currently, the tools to create scaffolds and structures for organ regeneration are limited, and to have precise control over the mechanical and microenvironmental variables is difficult. The other limitation is that there are only a few molecular, cellular, and tissue levels tests to evaluate what is actually happening within the constructs. However, tissue engineers are doing their best in trying to mimic the environment of the tissue. This section presents a brief account of the developments in several major areas of organ regeneration.

6.3.1 Prospects for liver regeneration Liver is one of the most perfused organs and is usually one of the first organs to be exposed to xenobiotics and pathogens. Liver cirrhosis, acute liver failure, and viral hepatitis are known to affect a significantly large number of patients. Many of them are hospitalized at critical stages of organ failure and need immediate organ substitution. Bioartificial liver support systems or liver bioreactors have thus occupied biomedical engineers’ efforts for a long time (Sussman and Kelly, 1993). Some of them are widely used clinically, while modifications on others continue to be made. They are, however, not permanent solutions. For such patients, liver tissue engineering is an exciting prospect. Fortunately, compared to functional cells of other vital organs, liver cells (hepatocytes) can regulate their growth and size in response to damage. This phenomenon is referred to as compensatory hyperplasia (Kay and Fausto, 1997). Thus, obtaining part of liver cells for allogenic or autologous engineering does not cause severe damage to the donor organ. However, obtaining donors for all such patients is a daunting task, as indicated by the American Liver Foundation – only 3000 donors for about 30 000 patients affected with one form or another of liver failure

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(Langer and Vacanti, 1993). In such cases the patients may be assisted on an artificial bioreactor, and meanwhile a fully functional organ can be grown from hepatocytes of the donor or the patient himself. Functional tissue engineered liver can then be transplanted. It has been possible to derive functional liver cells from human embryonic stem cells. Also hepatocyte precursor cells have been isolated successfully from liver tissue. With such advances it would be possible to grow the hepatocytes in vitro on supporting three-dimensional scaffolds, generating artificial livers ready for transplantation. A second avenue for liver tissue engineering lies in treatment of metabolic disorders by gene therapy (Kay and Woo, 1994). Liver is a major biosynthetic factory of the body and is responsible for secretion (and removal) of many essential biomolecules in (from) circulation. Tissue engineering promises cures for various hepato-deficiency disorders that manifest in metabolic abnormalities. The methodology adopted involves isolation of the hepatocytes incapable of synthesizing a particular molecule due to genetic abnormality, followed by genetic engineering of these cells, and their re-introduction in patients. Such attempts have been tried in patients suffering from familial hypercholesterolamia (Grossman et al., 1994).

6.3.2 Myocardial tissue engineering Worldwide, thousands of patients are afflicted with myocardial infarction (MI), which deteriorates to advanced heart failure despite optimal medical therapy. MI results in more hospitalization than all forms of cancer and is associated with a significant number of morbidities (Capi and Gepstein, 2006; Jessup and Brozena, 2003). Limitations of current therapies and transplantations have pushed great interest in tissue engineering to repair the failing heart. Human heart comprises terminally differentiated cardiomyocytes, which have very limited potential for self-renewal. Today, unlike heart valves or blood vessels, heart muscle has no replacement alternatives and engineering a heart muscle remains the most critical factor for myocardial tissue engineering. Towards this end, various cell sources have been investigated for their potential to yield functional cardiac muscle cells. Fetal cardiomyocytes, embryonic stem cells, skeletal myoblasts, crude bonemarrow cells, mesenchymal stem cells, hematopoietic stem cells, fibroblasts smooth muscle cells, etc., have been proposed by various authors. Recent reports indicate that mesenchymal stem cells collected from the menstrual blood are the best sources for cardiomyocyte differentiation. Extensive work in the isolation of correct cell type is still required, as the ability of skeletal cells to reproducibly differentiate into contractile and electrophysiologically functional cardiac cells is debated, and concerns over their safetyrelated issues (arrhythmias for skeletal progenitors and calcification with

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bone marrow cells) remain unaddressed (Leor et al., 2005). For in vitro cardiac tissue engineering, bioreactors are designed to simulate adjustable pulsatile fluid flow and varying levels of pressure, as encountered by cardiac cells in real time situations. Investigations by one group have identified deciding factors which lead to strongly contracting (up to 3 mN/mm2) and morphologically highly differentiated cardiac muscle constructs named ‘engineered heart tissue’ (EHT). They believe that (i) addition of Matrigel to the reconstitution mixture (only in rat EHT), (ii) EHT culture under cyclic exposure to mechanical loads, (iii) circular shape of the bioreactor, in contrast to EHT patches, and (iv) utilization of cell mixtures rather than purified cardiac myocyte populations are key factors for the in vitro construction of EHT (Zimmermann et al., 2004). The achievements in the field of myocardial tissue engineering are applied in parallel to the repair of congenital cardiac defects. For example, a study has demonstrated full replacement of the ventricular free wall by seeding mesenchymal stem cells in a scaffold made of polytetrafluoroethylene, polylactide mesh, and Type-I and -IV collagen hydrogel in an animal model is possible (Krupnick et al., 2002).

6.3.3 Corneal regeneration and tissue engineering of the eye Of various ocular problems, corneal and conjunctival epithelial cell injury, degenerations, and abnormalities are relatively common. Persistent epithelial defects caused by microbes, chemicals, iatrogenic, physical, and congenital insults, along with retinal deficiencies pose strong threats to vision. Surface diseases such as Stevens–Johnson’s syndrome, chemical and thermal burns, recurrent pterygia, ocular tumors, immunologic conditions, radiation injury, inherited and congenital syndromes, aniridia, and ocular pemphigoid, severely compromise the ocular surface and cause catastrophic visual loss in otherwise potentially healthy eyes. Treatment is expensive, frustrating, time-consuming, and often unsuccessful (Schwab, 1999). Tissue engineering of autologous limbal epithelial cells cultured on amniotic membrane and 3T3 fibroblasts is found to be a simple and effective method of reconstructing the corneal surface and restoring useful vision in patients with unilateral deficiency of limbal epithelial cells (Pellegrini et al., 1997; Tsai et al., 2000). This is also discussed in more detail in an earlier chapter on this topic in this book.

6.3.4 Cartilage tissue engineering Damage to the articular cartilage by trauma or degenerative joint diseases, such as primary osteoarthritis, causes disabling joint pains. Articular

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cartilage facilitates movement by minimizing friction between joints (as lubricants) and allows load-bearing through distribution of mechanical stress and acts as a shock absorber against compressive forces. Once damaged, cartilage tissue possesses very limited potential for healing. Current treatment methods for restoration of function of articular cartilage, other than total joint arthroplasty, are autografting, allografting, periosteal and perichondrial grafting, stimulation of intrinsic regeneration by intentionally drilling full-thickness defects, pharmacological intervention, and, autologous cell transplantation such as the periosteal flap technique marketed by Genzyme Corp. (Cambridge, MA, USA). Despite these, cartilage damage often cannot be repaired to a fully functional normal state (Brittberg et al., 1994; Tuan et al., 2003). Scaffolds such as alginate and collagen hydrogels which can maintain the rounded morphology of chondrocytes are some of the favorable cell carriers for cartilage tissue engineering (Kuo and Tuan, 2003). In addition, genetic elements responsible for chondrogenic differentiation have been identified and will aid articular engineering (Bursell et al., 2007).

6.3.5 Tissue engineering of the bone As life expectancy increases, so does the need to treat large-sized bone defects. New biomaterials combined with osteogenic cells are now being developed as an alternative to bone grafts. For example, bioceramics such as hydroxyapatite and tricalcium phosphate can be manipulated to match the mechanical properties of bone and are generally considered osteoinductive (inducing new bone formation), and with highly porous architecture, they are considered osteoconductive (favoring new bone tissue ingrowth). Porous ceramics of hydroxyapatite and β-tricalcium phosphate loaded with MSCs were shown to be capable of healing critical-sized segmental bone defects not capable of being healed by implantation of scaffold alone. MSCs seeded in bioceramics have been shown to regenerate bone in a variety of studies (Bruder et al., 1998). Bone tissue formation after in vivo transplantation of autologous bone marrow-derived cells in the bone defect of a geriatric patient has been achieved in the clinic. Bone marrow cells were expanded in vitro during which time they were exposed to novel recombinant human transforming growth factor 1, fusion protein bearing a collagen-binding domain, dexamethasone and β-glycerophosphate, followed by loading them into porous ceramic scaffolds (Becerra et al., 2006). Current research is focused on evaluation of biomaterials for their ability to promote osteogenesis (Hee and Nicoll, 2006; Jager et al., 2005).

6.3.6 Vascular and valvular tissue engineering Peripheral vascular diseases and diseases of coronary arteries, e.g. atherosclerosis, have immense implications on healthcare. Coronary artery bypass

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and balloon angioplasty, along with use of pharmacotherapy are current treatment modalities. While balloon angioplasty is considered a simple and effective tool in the armory of interventional cardiologists, recurrent restenosis largely limits its universal application. Working out a solution to recurrent restenosis after balloon angioplasty of arteries is an exciting area of research in itself and has led to the development of drug-eluting stents (King, 2007). Autologous coronary bypass surgery, on the other hand, is difficult in a significant population of patients due to lack of arteries suitable for the purpose (Yow et al., 2006). Existing cardiac disease further complicates the condition of patients by retarding the natural healing process of both small capillaries (angiogenesis) and large vessels (arteriogenesis) (Hill et al., 2003). Thus, the great clinical need to restore vascular functions has driven the medical and bioengineering community towards fabrication of blood vessel substitutes, called ‘designer vessels’, tissueengineered vascular conduits, arterial grafts, etc. Such conduits must be nonthrombogenic, nonimmunogenic, and suturable, and must possess adequate mechanical strengths and appropriate functional and healing responses. Conduits made of Dacron or expanded polytetrafluoroethylene (ePTFE) are known to perform adequately in high-flow, large-diameter environments and were primitive strides in this area. However, these materials do not show proper utility for small-diameter vessel (<6 nm) defects, which are clinically more challenging to bypass (Berglund and Galis, 2003). Decellularized natural membranes seeded with endothelial progenitor cells are being evaluated to provide better vessel substitutes (Divya et al., 2007; Stosich et al., 2007). Heart disorders due to end-stage valvular failure and congenital valve deformities offer exciting challenges for a tissue engineer. They require complex mold design that can simulate the complex 3D geometry of trileaflet valves of the heart (Asnes et al., 2006). Trileaflet heart valve scaffolds fabricated from polyglycolic acid coated with poly-4hydroxybutyrate, sequentially seeded with autologous ovine myofibroblasts and endothelial cells, grown for 14 days in a pulse duplicator in vitro system under gradually increasing flow and pressure conditions, have been designed as bioartificial substitutes (Schmidt and Hoerstrup, 2005). Polyglycolic acid is also used for constructing vascular grafts with improved patency (Shao-jun et al., 2007).

6.3.7 Diabetic tissue engineering Diabetes affects 16 million people in the US, and it is estimated that in the near future India will be the diabetic capital of the world. Diabetes is of two types: the more common Type 2 or insulin independent/insulin resistant diabetes and the Type 1 or juvenile onset diabetes. In the last decade, changing lifestyles have increased the proportion of patients diagnosed with

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Type 1 diabetes. This disease is caused by autoimmune destruction of insulin-producing β islets in the pancreas. The resulting lack of insulin leads to hyperglycemia and its devastating secondary complications include cardiopathy, retinopathy, nephropathy, and neuropathy (Diabetes Control and Complications Clinical Trial Research Group, 1993). For tissue engineering based management, islet transplantation is preferred over whole organ transplantation. Islet transplantations are heterotropic; even if lodged in a site other than their normal anatomy, such as the liver, islet cells can still produce insulin and control blood sugar levels (Rossini et al., 1999). Islet transplantation received a new direction with the so-called Edmunton Protocol (Shapiro et al., 2000). However, the loss of viability and delicate nature of islet cells require extensive work on their culture conditions. Insulin expressing cells from mouse stem cells have been generated quite some time back (Lumelsky et al., 2001). Recent progress in pancreatic progenitor cells as a cell source and cell encapsulation, along with surface modification of islets for immune acceptance of islets, promise a good future for this global epidemic (Beck et al., 2007; Liao et al., 2007).

6.3.8 Neural tissue engineering Diseases of neuronal injuries are often difficult to cure by pharmacological interventions. These include traumatic injury to the brain and spinal cord, peripheral nerve damage resulting in paralysis, and the neurodegenerative trio of Parkinson’s, Alzheimer’s and Huntington’s diseases, affecting millions of patients worldwide. The blood–brain barrier (BBB), which regulates entry of chemicals into the brain, is a highly restrictive membrane and is impermeable to many drugs discovered for action against neurodegenerative diseases. In the case of spinal cord injuries, glial scarring due to accumulated astrocytes prevent nerve–nerve connections to be set in during the regeneration process. Similarly, surgical restoration of peripheral nerve connections results in significant donor site morbidity (Willerth and Sakiyama-Elbert, 2007). Though it is now well established that neurons have the capacity to renew themselves, the process is very slow. Typically, neuronal tissue engineering has focused on the development of nerve guidance channels or conduits that act as bridges for regenerating neurons. Conduits can be single or multiple lumens and incorporate neutrophic factors. Modulation of delivery kinetics of growth factors from conduits can be performed to accelerate nerve regeneration. Such an attempt with glial derived neurotrophic factor from conduits produced by spinning mandrel technology and consisting of a collagen tube coated with layers of poly(lactidecoglycolide) (PLGA) has yielded good results in animal models (Piquilloud et al., 2007). Fetal human mesenchepalic cells have been injected into patients with Parkinson’s disease and shown to differentiate into

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dopaminergic neurons, providing patient relief for up to ten years. However, variation in the viability of cells remains one cause of getting irreproducible results (Rosenthal, 1998).

6.3.9 Urethra tissue engineering Various biomaterials without cells, such as PGA and acellular collagenbased matrices from the small intestine and bladder, have been used experimentally (in animal models) for the regeneration of urethral tissue (Chen et al., 2000). Some of these biomaterials, such as acellular collagen matrices derived from bladder submucosa, have also been seeded with autologous cells for urethral reconstruction. Recently, scientists have been able to replace tubularized urethral segments with cell-seeded collagen matrices (Atala, 2002). Acellular collagen matrices derived from bladder submucosa have now been used experimentally and clinically. In animal studies, segments of the urethra were resected and replaced with acellular matrix grafts in an onlay fashion. Histological examination showed complete epithelialization and progressive vessel and muscle infiltration, and the animals were able to void through the neo-urethras. These results were confirmed in a clinical study of patients with hypospadias and urethral stricture disease. Decellularized cadaveric bladder submucosa was used as an onlay matrix for urethral repair in patients with stricture disease and hypospadias. Functional neo-urethras were noted in these patients with up to a seven-year follow-up. Unfortunately, the above techniques are not applicable for tubularized urethral repairs. If a tubularized repair is needed, the collagen matrices should be seeded with autologous cells to avoid the risk of stricture formation and poor tissue development (Atala, 2002). Therefore, tubularized collagen matrices seeded with autologous cells can be used successfully for total penile urethra replacement. Whether bone marrow derived mesenchymal stem cells loaded on collagen matrix could help in such regeneration is not known. It would be also interesting to explore whether urethral linings have specific stem cells.

6.3.10 Engineering the bladder The success of cell transplantation strategies for bladder reconstruction depends on the ability to use donor tissue efficiently and to provide the right conditions for long term survival, differentiation, and growth. Urothelial and muscle cells can be expanded in vitro, seeded onto polymer scaffolds, and allowed to attach and form sheets of cells (Liu et al., 2009). In an elegant experiment carried out in animals, urothelial and muscle cells were separately expanded from an autologous bladder biopsy and seeded onto a bladder-shaped biodegradable polymer scaffold. The results from this

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study showed that it is possible to tissue engineer bladders that are anatomically and functionally normal. Clinical trials for the application of this technology are currently being conducted. These experiments are promising in terms of showing that engineered tissues can be implanted safely. Research is being carried out to accomplish the goal of engineering fullyfunctional bladders.

6.3.11 Genital tissues A wide variety of pathologic penile conditions, such as penile carcinoma, trauma, severe erectile dysfunction, and congenital conditions such as ambiguous genitalia, hypospadias, and epispadias, need reconstructive surgery. One of the major limitations of phallic reconstructive surgery is the scarcity of sufficient autologous tissue. The major components of the phallus are corporal smooth muscle and endothelial cells. The creation of autologous functional and structural corporal tissue de novo would be beneficial. Derivation and identification and ex vivo expansion of autologous mesenchymal stem cells from these tissues would be highly beneficial for such reconstructive surgeries. In recent studies, autologous cavernosal smooth muscle and endothelial cells were harvested, expanded, and seeded on acellular collagen matrices and implanted in a rabbit model (Kershen et al., 2002; Kwon et al., 2002). Histologic examination confirmed the appropriate organization of penile tissue phenotypes, and structural and functional studies, including cavernosography, cavernosometry, and mating studies, demonstrated that it is possible to engineer autologous functional penile tissue. It will be interesting to see if mesenchymal stem cells along with other stem cell sources such as the endothelial progenitor cells from blood could be used together to creat the implants for therapy. Congenital malformations of the uterus may have profound implications clinically. Similarly, several pathologic conditions, including congenital malformations and malignancy, can adversely affect normal vaginal development or anatomy. Patients with cloacal exstrophy and intersex disorders may not have sufficient uterine tissue present for future reproduction. Tissue engineers have started to investigate the possibility of engineering functional uterine tissue using autologous cells. Autologous rabbit uterine smooth muscle and epithelial cells were harvested, then grown and expanded in culture. These cells were seeded onto preconfigured uterine-shaped biodegradable polymer scaffolds, which were then used for subtotal uterine tissue replacement in the corresponding autologous animals. Upon retrieval six months after implantation, histological, immunocytochemical, and Western blot analyses confirmed the presence of normal uterine tissue components. Biomechanical analyses and organ bath studies showed that the functional characteristics of these tissues were similar to those of normal

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uterine tissue. Vaginal reconstruction has traditionally been challenging due to the paucity of available native tissue. Menstrual blood mesenchymal stem cells have been shown to have a propensity to differentiate into myometrial cells; also, now it is possible to collect, isolate and expand the mesenchymal stem cells from endometrial biopsies. These advances and findings indicate that a combined approach of using various stem cells and a biomaterial scaffold may turn out to be the easiest way to reconstruct vaginal tissue. Such in vitro vaginal models would also be interesting in developing in vitro embryonic implantation models to study maternal fetal interactions and use them as drug screening platforms. Efforts towards developing such in vitro models is currently being carried out in our laboratory, by employing various endometrial and other stem cell sources.

6.4

Conclusions

Clinical evaluation of tissue engineered implants suggests a modest impact on healthcare so far. Modern state-of-the-art in stem cell biology, cell culture methods, molecular genetics and their application to developmental biology and pharmacology, including animal models, and development of biomaterial science, are set to deliver on the expectations harbored on tissue engineering. Although stem cells have been known for some time, the biology of stem cells and their manipulation for therapeutic purposes have become the subject of intense research only in the last decade. The right source of stem cells for organ regeneration will continue to be a debate for the coming years. Recent developments in the field of biomaterials and tissue engineering as described in all the previous chapters in this book indicate a quantum leap of this field over the last decade. It can be envisioned that, once fully developed, tissue engineering will radically change the way we look at diseases and aging. However, to protect from misuse of this powerful technology, carefully drafted ethical regulations will be required.

6.5

References

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7 Materials facilitating protein drug delivery and vascularisation P. M A RT E N S, A. N I L A S A R OYA and L. A. P O O L E-WA R R E N, University of New South Wales, Australia

Abstract: Hydrogels are versatile materials that have many applications including drug delivery and tissue engineering. Individually, synthetic and natural hydrogels have many advantageous properties; however, their disadvantages have resulted in an increased research focus on biosynthetic hybrids. The key for many biomedical applications is the ability to produce vascularised structures which allow transport of nutrients and waste products. Critical challenges in producing such biosynthetic hydrogel constructs for tissue replacement are in selection of the appropriate biological signals and incorporation without loss of function. Overcoming the normal wound-healing response and producing physiological vascularisation is an ongoing research area. Key words: hydrogels, bio-synthetic matrices, protein delivery, vascularisation.

7.1

Introduction

Hydrogels are networks of insoluble polymer chains that have a high weight percent of water. As a result of this high water content and the polymer chain interactions within the network, the structure of synthetic hydrogels has been compared with that of the extracellular matrix (ECM) of tissue.1 This attractive feature of hydrogels has been exploited to produce materials that either allow simple delivery of soluble bioactive factors or, at the other end of the spectrum, present insoluble biological molecules to modulate cell and tissue functions. Delivery of biological signals such as growth factors or ECM is a key requirement for many medical applications today and the specific approach used depends largely on the application site. This chapter will examine the application of composite hydrogels for both cell and drug delivery, with emphasis on the mode of presentation of biological components and the use of these constructs for enhancing vascularisation. Figure 7.1 shows a schematic that summarises the major ways that biological information can be incorporated into hydrogels, from entrapment 179 © Woodhead Publishing Limited, 2010

(a)

(b)

(c)

7.1 Schematic of the various ways biological proteins/polymers can be incorporated into synthetic hydrogel networks. (a) The biological protein is simply mixed in with the network and is free to diffuse out; (b) the biological protein is chemically bound to the network and is present on the surface; (c) a combination of (a) and (b), where one type of protein is bound to the network and remains with it, while another protein is mixed in and can diffuse out. © Woodhead Publishing Limited, 2010

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through to crosslinked or bound molecules, to combinations of these approaches.

7.1.1 Protein delivery – presentation or release? Hydrogel systems can be designed to deliver soluble bioactive factors such as hormones and growth factors, either via incorporation of the active agent and simple passive diffusion or through the design of the hydrogel system to specifically interact with the bioactive agent and thus sequester it for ‘release on demand’. Systems that rely on passive diffusion typically involve entrapment of bioactive molecules prior to hydrogel formation2 or uptake via swelling of hydrogels in the presence of the bioactive of choice.3 Both approaches can be used to incorporate high molecular weight drugs into hydrogels, with subsequent release kinetics usually being described by Fickian diffusion. The more complex systems that sequester bioactive factors typically include as components biological molecules that interact specifically with, or are covalently bound to, the active agent to be delivered.4,5 This latter approach may also enable more physiological presentation of the soluble factor such as in heparin presentation of basic fibroblast growth factor.4,6 In these types of hydrogels, a biopolymer, in this case the highly sulphated polysaccharide heparin, is incorporated as a component of the hydrogel and its growth factor binding function is exploited as a method for controlled release. Similarly, systems that rely on enzymatic release of growth factors from fibrin matrices have been engineered.5 One example of these types of systems used a vascular endothelial growth factor isoform modified with a factor XIIIa substrate site from α-plasmin inhibitor, allowing it to specifically bind with the fibrin hydrogel delivery vehicle. Cells invading the fibrin matrix degraded the matrix and thus released growth factor at a rate equivalent to cellular ingrowth.5 It is clear that hydrogels offer significant flexibility in design and, due to their inherent similarity to the ECM of tissue, are a useful base for delivery and presentation of biological information to allow cell interactions. In order to understand the variety and flexibility available in hydrogel systems, it is important to first define the classes of hydrogels that have been described.

7.2

Hydrogel classification

There are many different approaches to the classification of hydrogels, and these can be based on chemistry, structure, crosslinking method or end use. For the purposes of this chapter, hydrogels will be put into one of three categories: natural, synthetic or composite bio-synthetic. Each of these types of polymers has its own benefits and drawbacks (see Table 7.1), which

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Table 7.1 Outline of some of the key benefits and drawbacks for natural and synthetic polymeric hydrogels

Natural

Advantages

Disadvantages

Examples

•

•

Batch to batch variation Supply and cost Immunogenic issues Possible prion transfer Usually mechanically weak

• Collagen • Elastin • Silk proteins • Alginate • Glycosaminoglycans • Polysaccharides

No biological recognition Potential toxic by-products from synthesis

•

•

Biological markers Biological signalling

• • • •

Synthetic

• • •

•

Bio-synthetic

•

•

•

Relatively cheap Easy to modify Consistency between batches Engineered variations in mechanical strength and network structure

•

Biological markers and signalling Dimensional stability from the synthetic polymer Engineered variations in mechanical strength and network structure

•

•

• • •

•

Same as above, but hopefully in a reduced form

• • •

Poly(vinyl alcohol) (PVA) Poly(ethylene glycol) (PEG) Poly(acrylic acid) (PAA) Poly(methyl methacrylate) (PMMA) Poly(hydroxy ethyl methacrylate) (pHEMA) PVA/heparin PEG/heparin PEG/VEGF

may make them more or less appropriate for a particular application. Synthetic and natural polymer hydrogels crosslinked via a number of different methods have long been studied as delivery vehicles for high molecular weight drugs.2,3,7 More recently, increasingly complex systems delivering drugs and genes have been reported.8,9 A brief overview of these hydrogel systems is provided below.

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7.2.1 Natural hydrogels Natural hydrogels are defined as hydrogels that are made up of polymers that originate from living systems, such as humans, animals or plants. They include proteins, polysaccharides and nucleic acids such as deoxyribonucleic acid (DNA). Collagen is probably the most well known and researched natural hydrogel system, as it is the main constituent of most ECM in the human body.10 Once collagen is isolated from tissues and purified, it can be easily made into a natural hydrogel through exposure to temperatures of 37 ˚C. Collagen, like most natural polymers, has an inherent structure or conformation. When the appropriate conditions, such as exposure to body temperature equivalents, are presented, these polymers revert back to their native configuration which results in physical crosslinking between the chains. These physical crosslinks are often reversible and this may lead to rapid and unpredictable degradation in vivo. In addition, natural hydrogel systems extracted from native matrix often have poor mechanical properties and dimensional stability, are expensive and difficult to source and are potentially immunogenic. Attempts to improve the mechanical properties of such reconstituted natural hydrogels have been made through the introduction of chemical crosslinkers (e.g. glutaraldehyde), although these chemicals often have detrimental effects on incorporated cells, proteins or drugs.11 More recently, recombinant collagens of various types have been developed and these have potential to overcome some of the aforementioned shortcomings.12 Elastin, another ECM protein has also been produced in hydrogel scaffold forms and may prove an alternative or adjunct to collagen in applications where greater elasticity is required.13,14 Several authors have reported use of isolated elastin fibres for tissue engineering; however, soluble elastin, in particular when formed in a blend with Type I collagen, may promote development of new blood vessels in vivo without calcification occurring.15 Although few studies have been reported on pure elastin hydrogels, gels formed via coacervation of elastin-like peptides have been shown to support survival and differentiation of encapsulated cells.16 More recently, human recombinant tropoelastin has been produced as a hydrogel under alkaline conditions in the absence of crosslinking agents.17 These studies demonstrated that dermal fibroblasts proliferated on the formed hydrogels, although encapsulation studies were not conducted. Thus, although natural polymers have some disadvantages relating to their physical and mechanical properties, due to their biological design, interactions with cells and tissues are typically excellent and it is these properties that are not found in synthetic polymers.

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7.2.2 Synthetic hydrogels Synthetic hydrogels are made from polymers that are manufactured in the laboratory and are not sourced from living tissues. These are the most common type of hydrogels, and there is a plethora of synthetic polymers available. Synthetic hydrogel systems are made from either polymeric chains that are subsequently crosslinked together, or from monomers via conventional polymer synthesis. In either case, the resulting network is a crosslinked, hydrophilic system. The benefit of synthetic hydrogels over natural hydrogels is that there is often more control over their structure and thus many of the network properties are more easily engineered. Modification of the molecular weight of the polymers, the amount of polymer in the system and the extent of crosslinking, is easily achieved using a variety of synthesis and fabrication approaches. This ability to manipulate the polymer characteristics usually results in greater tailorability of the resulting hydrogel’s properties, such as swelling, mechanics, mesh size, and degradation, and thus can result in better control of protein or drug release.18–20 This type of simple system is illustrated in Fig. 7.1a. Some of the most common synthetic hydrogel systems include poly(vinyl alcohol) (PVA), poly(2-hydroxyethyl methacrylate) (polyHEMA), poly(acrylic acid) (PAA) and poly(ethylene oxide/glycol) (PEO/PEG). These polymers can be made either directly into hydrogels via physical crosslinking methods or through the addition of crosslinker molecules, or can be further functionalised with crosslinkable groups and polymerised into a hydrogel network. As noted above, synthetic polymer hydrogels can be useful as drug delivery matrices where cell interaction is not required; however, they do not typically support attachment, proliferation and differentiation of living cells.

7.2.3 Composite bio-synthetic hydrogels The incorporation of biological polymers, such as proteins and polysaccharides, within a synthetic matrix to form a composite bio-synthetic hydrogel system is another approach to delivering biological information to support cell growth and differentiation.6,21 This design approach forms the basis for application of hydrogels to cell encapsulation and tissue engineering matrices. The synthetic polymer is able to provide dimensional stability and appropriate mechanical properties while the biological polymer presents appropriate ligands for cell attachment, presents cytokines such as growth factors, or initiates other critical cell functions. Figure 7.1b shows a schematic of the major functional components of such systems. Combination of both the capacity for delivery of high molecular weight drugs and incorporation of one or more biological polymer types into a

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synthetic matrix makes bio-synthetic hydrogels attractive alternatives to ‘bulk’ polymer scaffolds. These types of systems clearly have properties most similar to natural ECM due to the incorporation of biological moieties for cell interaction, as well as the capacity to present soluble factors via these structures and the ability of the active factors to diffuse throughout the matrix at a similar rate to that in ECM. Figure 7.1c presents schematically the variety of functions these latter systems can potentially fulfil. There are several methods of forming the bio-synthetic composites, each with their own advantages and disadvantages. The synthetic and biological polymer can be either simply mixed together forming an interpenetrating network (IPN), the synthetic polymer can be crosslinked and the biological polymer can be simply blended with the system (a semi-IPN), or both types of polymers can be crosslinked together. Alternatively, biological signals can be covalently attached to the synthetic polymer backbone, which can then be crosslinked to form a network displaying the biological information. The choice of which type of network to use depends on the polymer properties and the end application of the system. In many cases, the properties of one or both of the polymers will change when used in conjunction with the other polymer, and this may have an impact on solubility, biological function, and crosslinking.22 Regardless of the hydrogel platform used, or the approach used for incorporation of biological signals, careful characterisation of polymers before and after modification and incorporation into the hydrogel are essential. This allows identification of the mechanisms of protein or other drug release, and presentation of signals, and aids prediction of how hydrogel design may impact on these factors.

7.3

Factors influencing protein encapsulation and release

Ideally, the release of proteins and drugs from a scaffold in vivo will occur over an extended time period, which allows for a greater effective lifetime of the drugs. The ability to control and predict the release is of utmost importance and there are many factors that influence protein encapsulation and release rate. Passive diffusion is the simplest and most common form of release exploited in hydrogel drug delivery systems (see Fig. 7.1a). Many excellent reviews highlighting the range of backbone chemistries and crosslinking approaches used, as well as the use of pre-formed or in situ forming hydrogels have been published and can be referred to for more information on this subject.1,23,24 In systems where the protein is simply physically blended with the hydrogel system, the release of protein is usually controlled by simple diffusion

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through the network or through combined degradation of the network and diffusion. In both of these cases, the physical structure of the network will have a large impact on release. Research has focused on slowing down the release/diffusion of protein based drugs through changes to the hydrogels’ network structure, as well as via alteration of interactions between the polymer and the proteins. Changing the network structure is relatively straightforward and has important implications on the diffusion of proteins from the hydrogel. However, controlling the interactions between the hydrogel and the proteins is more complicated and can be achieved via a range of approaches. The chemistry of the backbone polymer can not only influence the overall structure, but can also influence interactions between protein and synthetic polymer chains. Often the chemistry is chosen such that partitioning or complexation of the protein to the polymer backbone will occur, thus altering the release profile. In addition, the polymer and the protein can be chemically conjugated to each other, which will have profound effects on the rate of release.

7.3.1 Network structure Alterations in the network structure of hydrogels may be one of the most straightforward methods of varying the release profile of encapsulated proteins. The effective area for protein diffusion is characterised by the average mesh size of the network (ξ), and mesh size can be broadly defined as ‘the space between macromolecular chains in a crosslinked network and is characterised by the distance between two adjacent crosslinks’.25 If the encapsulated proteins are smaller than the mesh size of the hydrogel network and do not have any interactions with the network, then Fickian diffusion will control their release. In this case, simple changes to the network structure will result in changes to the release profile.26 Since the mesh size is directly related to the average molecular weight ——— between crosslinks (Mc),25,27 one of the simplest approaches to modifying network structure is to change the number or spacing of the crosslinks in the system, as shown in the schematic in Fig. 7.2. For some systems, such as collagen and gelatin, the concentration of chemical crosslinker such as glutaraldehyde added during the formation of the network will affect the crosslink density.28,29 In other systems, such as polyvinyl alcohol (PVA) and dextran, the degree of substitution (DS) (i.e. the amount of functional groups that are added to the polymer backbone) can be altered.29–31 The molecular weight of the backbone polymer chain can also be varied with the DS staying the same, which again results in a change to the mesh size.32 In PEG-based systems, the backbone polymer and the crosslinking molecule are the same, so an increase in PEG chain length results in an increase in the mesh size.27 As would be expected, if all other factors remain constant,

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(a) Mc Mc

(b)

x

(c)

7.2 Schematic of two possible ways of decreasing the mesh size (ξ) of hydrogel systems. (a) The original network, (b) the molecular weight of the backbone polymer chain remains the same and the number of crosslinks has been increased, and (c) the molecular weight of the backbone polymer chain has been decreased and the number of crosslinks remains the same.

and the DS is increased, the network mesh size must decrease, which will lead to increased hindrance to any diffusing molecules.30 Another way of altering the mesh size or network structure is to influence the amount of cyclisation that is occurring. An increase in cyclisation means a decrease in the amount of effective crosslinking groups, which results in an increase in the mesh size of the system. Increases in cyclisation can be achieved through the addition of more water at the time of polymerisation.29 Cadee et al. have utilised this network feature and have polymerised a series of dextran based gels with a range of initial water contents. Gels with high initial water content (and thus larger mesh sizes) were shown to have Fickian release of recombinant interleukin-2 molecules, whereas no release was observed for gels with low initial water contents.30 Thus, by simply decreasing the amount of water present at the time of

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polymerisation, the overall release amount and release rate can be significantly altered. In all of these cases, if the network structure is altered such that the molecular weight between crosslinks is increased, the average mesh size of the system will be increased. A larger mesh size will result in easier diffusion of the protein chains out of the system. Conversely, if the chains are diffusing too quickly, the mesh size and network structure can be altered such that it hinders the diffusion of the large molecules (see Fig. 7.2).

7.3.2 Polymer chemistry Changing the chemistry of the polymer backbone can not only effect the interactions between the polymers and the proteins (see Section 7.3.3), but can also influence the mesh size and the release kinetics. The chemical nature of the side groups that are attached to the polymer chain can have a strong influence on the network structure and can often be picked to respond to their environmental conditions, such as pH, temperature and ionic strength. These gels are sensitive to their environment and are generally collapsed with a small mesh size in one environment and swollen with larger mesh sizes that allow protein diffusion in another environment. There are many examples of stimuli-sensitive hydrogels,33–36 and just a few brief examples are given here. Poly(methyl methacrylate) (PMMA) has been shown to either collapse or expand on a PVA sheet, depending upon the pH of the environment. If the PMMA is expanded, it blocks the pores and the permeability of the system decreases, whereas if the PMMA is collapsed on itself, the pores remain open and diffusion is increased.37 Similar results were found with polyacrylic acid/PVA hydrogels, where a change in pH resulted in changes to the hydrogen bonding between the polymers and the network mesh size.38 These hydrogels have the significant advantage of being able to be loaded with the drugs/proteins outside of the body in the swollen state, and then collapsed to entrap the proteins until they reach the site within the body where the protein is required. One drawback of these systems is that, generally, once they are in the swollen state, the proteins readily diffuse out of the system. Thus a longer sustained release is not achieved. Another part of the polymer’s chemistry that can be altered to affect the release of encapsulated proteins is the incorporation of degradable linkages. By making the hydrogel degradable, the release of the protein can be changed from being diffusion controlled to degradation controlled. To limit the diffusion of the protein, and to control it by degradation, the initial mesh size of the network should be smaller than the hydrodynamic radius of the protein. In this manner, the protein cannot diffuse through the network until degradation occurs to an extent such that the average mesh

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size matches the size of the protein. Depending on the type of degradation and relative speed of diffusion once the degradation has occurred, the release can be either entirely controlled through the degradation or through degradation and subsequent diffusion.22 Similar to the network structure, for proteins and drugs that are simply mixed into the polymer system (i.e. non-covalently bound), an increase in the extent of crosslinking leads to a decrease in the degradation rate. Therefore, if the release of the proteins depends on degradation, increasing the level of crosslinking will increase the retention time and decrease the protein release rate. Tabata et al. used glutaraldehyde to crosslink collagen gels, and increasing the amount of glutaraldehyde resulted in an increase in the crosslink density which also resulted in slower in vivo degradation.28 These researchers then incorporated vascular endothelial growth factor (VEGF) into the collagen hydrogels and a direct correlation between VEGF release and hydrogel degradation could be observed. Increasing crosslinking density resulting in slower degradation and slower VEGF release.28 The release can also depend on both diffusion and degradation, as was demonstrated for Pluronic gels.39 The surface layer of these gels experienced an increase in swelling, which leads to the diffusion of the proteins. However, the gels are also being degraded and the mass erosion of the Pluronic gels was shown to correlate very closely with release of encapsulated human growth hormone.39 More advanced degradable systems can also be developed, such as the incorporation of protein-loaded microspheres into a hydrogel system. The particles and the hydrogel will thus work in parallel, with the proteins having to be released from the spheres and then also from the gel. If the protein is wholly contained within the spheres, this overall hydrogel structure should lead to a lower initial burst and a more sustained release. In addition, the network structure and crosslinking density of both phases can be controlled to allow for greater control over the release profile.40

7.3.3 Partitioning and complexation Altering the hydrogel system to increase the interactions between the polymer chains and the releasing molecules is another popular research area. An increase in polymer–protein interaction can be achieved through the proper choice of backbone chemistry. Positively charged hydrogels, such as gelatin, will interact and bind with negatively charged proteins (VEGF) and DNA.41 This increased interaction will result in a slower release. In addition, positively charged sidegroups could also be chemically bonded onto the polymer backbone, which will again allow for the control of the release profile of the encapsulated proteins.42 If the interaction is strong enough, the molecule could be bound so tightly to the polymer network that the only way to release it is to degrade the network. Kushibiki et al.

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observed that for cationised gelatin hydrogels that had incorporated plasmid DNA, the release of the DNA was directly correlated to the gelatin degradation.42 Similarly, hydrophobic side groups can also be bonded to the hydrophilic polymer backbone, which allows for interactions with proteins.43 Leonard et al. covalently bonded alkyl chains to alginate polymers and used these amphiphilic systems to encapsulate a series of proteins (e.g. BSA). Again, the interaction between the polymer network and the protein molecules was shown to be strong enough to dictate that the release of BSA was directly related to the degradation of the alginate, and limited the diffusional release of BSA.43

7.3.4 Polymer–protein conjugates In the previous sections (7.3.1–7.3.3) it was assumed that the protein molecules to be delivered were simply mixed/encapsulated into the delivery system. However, proteins can be directly and covalently linked to the polymer delivery system. In this way, the protein/polymer system can be either non-degradable and simply present the protein,6 or the system can be degradable, where the protein release will be controlled via the degradation of the polymer.44 The conjugation of polymers and proteins has many advantages and disadvantages. The main advantage of conjugation is that conjugated proteins should have longer circulation half-lives as compared to non-conjugated proteins,45 as well as influencing the solubility46 and proteolysis susceptibility of the proteins.47 However, the process of chemically modifying the proteins may lead to alteration of the proteins’ biological activity. Thus, careful selection of how and where the conjugation will take place is of utmost importance if the benefits of the protein are to be preserved.6 For degradable polymer–protein systems, there are two main methods of protein release. The polymer system can degrade, and the protein is thus released in a conjugated form, or the linkage between the polymer and the protein can degrade.22 For example, Seliktar et al. studied the release of VEGF from multi-arm PEG hydrogels, where one case had VEGF bound by a non-degradable linkage and in the other the VEGF was bound by a degradable linkage. In both cases, the base polymer was degradable, but the overall release of VEGF was determined by how it was attached to the base network.48 When the proteins are bound to the polymer chains through nondegradable linkages, new molecules (often referred to as prodrugs) are formed. The release of the prodrugs is dependent on the degradation profile of the base polymer, and generally the prodrug is a biologically inactive form of the drug which becomes active through a transformation within the body.49

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A very common method of conjugation is the PEGylation of proteins. The addition of PEG to the proteins can make them more soluble, and also increase the effective molecular weight of the protein.46 These PEGylated proteins can be covalently bound to the surface of a base hydrogel or encapsulated into the hydrogel system for delivery. If the proteins are bound to the surface of a non-degradable hydrogel, they will simply be present in the system and perform their required function (i.e. the signalling of cells to adhere and proliferate).50 If either the proteins themselves (such as albumin, fibrinogen and collagen) or the base polymer are degradable, then the proteins will be slowly released into the body and the base polymer is used to control the structural and degradation properties.51 All of the factors/methods mentioned in this section are currently being studied to control the encapsulation and release of various proteins. There are numerous biomedical applications that would benefit from being able to reliably and accurately control where proteins are located and how/when they are released into the body. Tissue engineering is one such application, and in this case, the control of protein release is not only desirable it is absolutely essential.

7.4

Tissue engineering applications: vascularisation and protein delivery

Tissue engineering (TE) is a rapidly growing field and is a revolutionary approach for repair of damaged tissues and organs. TE involves the combination of a polymeric scaffold and encapsulated cells or biological factors, with the aim of repairing, replacing or regenerating lost or damaged tissue. Many different tissues are currently being studied, and some TE products have already been introduced to the market for skin and cartilage applications.52 One area of TE research that is currently receiving a lot of attention is the engineered delivery of growth factors in hydrogel scaffolds, with a particular emphasis on promoting blood vessel growth within the construct. In hydrogel TE constructs, both synthetic and natural polymers can be combined with cells and growth factors to produce complex structures. These types of structures are able to support viable cells and theoretically enhance development of tissue following implantation. As previously mentioned and outlined in Fig. 7.1c, biological polymer incorporated into the hydrogel network can act as a depot for growth factors, presenting them in a more physiological manner. For example, heparin copolymers with synthetic polymers have been demonstrated to bind and sequester basic fibroblast growth factor (bFGF) which maintains its function of enhancing cell adhesion, proliferation and differentiation.6,53 This approach may also

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be exploited to enhance vascularisation of TE constructs based on hydrogels, a critical limiting factor in engineering of large volume tissue masses.

7.4.1 Vascularisation in physiological wound healing Development of vasculature is essential for maintaining perfusion of new tissues. In the context of human growth and development, the processes of angiogenesis and vasculogenesis are responsible for the ordered production of new blood vessels as tissues form. Vasculogenesis refers to the formation of vessels from endothelial precursor cells, whereas angiogenesis is the sprouting of new vessels from existing endothelial cells. These processes are critical in the progression of ordered wound healing for rapid development of new blood vessels to transport oxygen, nutrients and the cells involved in the repair process.54 Both of these forms of new vessel production may occur in an implanted TE construct. New blood vessel formation in normal wound healing is characterised by early migration of endothelial cells to the wound site. Rapid formation of capillary structures that make up a major component of granulation tissue in the healthy healing wound, typically proceeds over the first week post wounding. Concomitantly, deposition of ECM occurs and once this reaches an appropriate level, inhibition of angiogenesis occurs followed by regression of vascularisation.55 The newly formed wound tissue then becomes largely avascular fibrous tissue. In pathological angiogenesis, characteristic of tumours and other diseases such as arthritis, prolonged abnormal vessel growth occurs, driven by sustained production of angiogenic factors.55 The challenges for tissue engineering, particularly in light of the proposed use of a variety of angiogenic growth factors, are to maintain new blood vessels produced so that they can perfuse functional tissues, to prevent unregulated angiogenesis, and to minimise the fibrotic response. These objectives are the basis for much of the current research in this field.

7.4.2 Vascularisation of implanted constructs It has long been recognised that vascularisation is an essential component for survival of cells and tissues in vivo. Cells in most tissues are never more than 100 to 200 micrometres from capillaries, in order to overcome the diffusion limitations of oxygen, nutrients and metabolic waste products in the hydrogel-like extracellular matrix. Various approaches have been used in an attempt to create new vasculature in TE constructs. Fundamental design requirements are that the scaffold has the appropriate soluble and insoluble biological signals, adequate pore or mesh size, and tailored degradability to allow migration and proliferation of endothelial cells or vascular progenitor cells. A challenge associated with using bio-synthetic hydrogels in this

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context is that unless they are degradable, cell migration and tube formation within the gels is unlikely. Many studies have reported tube formation in collagen gels (see recent review by Davis and Senger56), but formation of vascular structures in synthetic hydrogel constructs is rare. Given that the above essential design requirements are fulfilled, vascularisation of constructs may be produced via a range of approaches, which are summarised in Table 7.2. These include loading and subsequent release of angiogenic growth factors for both in vitro and in vivo vascularisation, in vitro seeding of hydrogels with endothelial or progenitor cells, or development of prevascularised constructs in vivo. The discovery that growth factors alone can stimulate development of new blood vessels promoted broad experimentation using various cytokines such as vascular endothelial growth factor (VEGF), platelet derived growth factor (PDGF) and fibroblast growth factor (FGF). Pharmacological studies have been performed extensively on therapeutical administration of these factors to induce vessel formation and improve tissue perfusion.57,58 However, high doses or repeat injections are required to achieve significant response because of the short protein half lives and rapid diffusion out of target tissues, which also raises concerns about the potential side effects at distant sites. One of the most promising examples of vascularisation in a synthetic matrix is based on the belief that optimal blood vessel formation would be regulated by the synchronised work of various growth factors. Nonhydrogel matrices made from synthetic polymers such as poly(lacticco-glycolic acid) (PLGA) have previously been developed to deliver multiple growth factors in succession, and in these matrices initial growth factor delivery was controlled by the degradation rate of the bulk structure, while delayed release was achieved by pre-encapsulating another factor prior to scaffold processing.59 By first releasing VEGF, endothelial vessel formation was stimulated, and the subsequent release of PDGF initiated the recruitment of smooth muscle cells to stabilise the newly formed vessels. After four weeks subcutaneous implantation in rats, the dual-release scaffold was shown to induce higher vessel density and growth of more mature vessels compared to the responses from scaffolds containing either factor alone. Human microvascular endothelial cells (HMVECs) transplanted on VEGF-containing PLGA matrices were able to form human-derived vessels within seven days of implantation in mice; the combination of transplanted ECs and VEGF resulted in a significant increase in density of vessels within the scaffold.60 Most of the research on developing hydrogel matrices for angiogenic growth factor delivery has focused on localised and sustained delivery of the individual growth factors, with particular emphasis on improving the incorporation efficiency and controlling the release of these factors. The

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Table 7.2 Strategies for producing vascularised tissue engineered constructs Strategy

Description

Angiogenic growth factors

Growth factors (GFs) incorporated into 3D matrices for delivery. Rate of GF delivery controlled by: • Pre-encapsulation of GFs for sustained release and to protect from denaturation • Use of HS/heparin crosslinked matrices to increase affinity of GFs to the matrices for sustained release • Controlling the release of heparin, as co-factor in the delivery of GFs

In vitro seeding using endothelial or progenitor cells

In vivo vascularisation

Formation of vascular-like structures in hydrogel matrices derived from collagen or basement membrane proteins (Matrigel) Seeding of cells on matrices in vitro followed by immediate transplantation; using the matrices and host as bioreactors for vascular formation Implanted GF-loaded matrices encouraged vascularisation in tissues surrounding the implant (a) Multiple growth factor delivery from PLGA matrices for vessel maturation (b) Use of heparincrosslinked hydrogels (natural and synthetic) to enhance GF activity in promoting neovascularisation HS-crosslinked and GF-loaded collagen matrices encouraged sustained vascularisation throughout the hydrogel structure

References

Gu et al., 2004; Nakamura et al., 2008; Peters et al., 1998 Cai et al., 2005; Steffens et al., 2004; Wissink et al., 2001

Ishihara et al., 2003; Tan et al., 2008 Finkenzeller et al., 2007; Hughes, 1996; Korff et al., 2001

Melero-Martin et al., 2007; Peters et al., 2002

Richardson et al., 2001

Cai et al., 2005; Fujita et al., 2004; Tae et al., 2006; Yao et al., 2006

Pieper et al., 2002

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angiogenic potential of these growth factor-carrying matrices has mostly been evaluated in terms of their capability to support endothelial cell growth and to encourage vascularisation of surrounding tissue. Synthetic scaffolds have been loaded with natural hydrogel components such as alginate and heparin-sepharose beads. These natural polymers have the ability to interact ionically or specifically with the growth factors (see above).61 These interactions resulted in more sustained release of the growth factors from the matrix, while at the same time protected the factors from denaturation. When calcium-crosslinked alginate beads were used to preencapsulate VEGF prior to incorporation in PLGA matrices, the effect was not only an improvement in incorporation efficiency but also an observed increase in VEGF activity in stimulating endothelial growth in vitro.62,63 The affinity of growth factors to glycosaminoglycan (GAG) molecules, natural components of the ECM, has led to the development of growth factor release vehicles built of GAG chains, using synthetic polymers as hydrophilic crosslinkers. Heparin, in particular, provides an analogue of the heparan sulfate proteoglycans, molecules that have been known to regulate angiogenesis in vivo.64 Heparin-crosslinked collagen matrices have been shown to support the attachment and proliferation of human umbilical vein endothelial cells (HUVECs) in vitro.65,66 When HS-crosslinked collagen matrices were implanted, the presence of HS alone was shown to induce transient vascularisation at the periphery of the matrices, while pre-loading the matrix with bFGF was shown to encourage capillary infiltrations,67 and the networks remained vascularised during a 10-week implantation period. Similar matrices containing VEGF were also shown to increase the capillary formation in the chorioallantoic membrane of chicken embryo, and increased vascularisation of surrounding tissue after a 14-day implantation in rats.68,69 The formation of injectable heparin/synthetic polymer hydrogels for the sustained release of growth factors has also been studied.70,71 Heparin gels were formed by reacting thiol-functionalised heparin molecules with bifunctional PEG chains as crosslinkers. After a two-week subcutaneous implantation in mice, new microvessels were observed in the tissue surrounding the VEGF-loaded gel while almost no vascularisation was observed around the control heparin gel. However, no cells or vascular structures were found to infiltrate the gel. The crosslinked heparin was shown to be necessary to enhance the bioactivity of bFGF and in promoting neovascularisation; however, too high a heparin concentration was shown to result in high retention of the growth factor.70 Microparticles made of a composite of fragmin, a low molecular weight heparin, and protamine, have also been studied for FGF-2 delivery for up to two weeks, and neovascularisation and fibrous tissue formation was observed around the subcutaneous injection site on mice.72

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More recent potential designs have relied on the release kinetics of heparin to control the growth factor delivery. Microspheres made of alginate/hydroxyapatite composite have been investigated as delivery vehicles for heparin, with the inclusion of hydroxyapatite found to increase the encapsulation efficiency but also increased the rate of heparin release.73 The use of heparin as a co-release agent to growth factors has also been investigated in chitosan hydrogels. When FGF-1 and FGF-2 incorporated chitosan hydrogels were implanted into the back of a mouse, neovascularisation was observed around the implant and this effect was increased when heparin was co-incorporated with the FGF.74 Most of the growth factors (∼75%) remained in the chitosan hydrogels at equilibrium, indicating that the growth factor release in vivo was done by degradation of the chitosan network. Instead of using GFs to encourage the infiltration of native cells into the matrix, other research groups have looked at directly incorporating isolated cells into the matrix before implantation. Loading scaffolds using cells with potential to produce vascular structures and culturing in appropriate bioreactor conditions can theoretically allow a ‘vascularised’ structure to be produced in vitro.75 Tube formation by endothelial cells appears to develop most efficiently in collagen and other natural polymer gels such as Matrigel,56,75,76 and co-culture with other cell types such as fibroblasts has been proposed as stimulatory to the process.77 Direct co-culture of smooth muscle cells (SMCs) with endothelial cells has demonstrated that SMCs induced endothelial cell quiescence and controlled the responsiveness of endothelial cells to angiogenic factors.78 The phenotype of SMCs embedded in cylindrical collagen modules was also found to influence the phenotype of HUVECs seeded on the surface of the modules, a response that can be modulated by adjusting the serum concentration in the SMC culture medium.79 However, successful development and implantation of in vitro grown constructs has not yet been reported. A variation on this approach is to seed constructs with progenitor cells in vitro and implant them immediately, theoretically using the recipient as a ‘bioreactor’ to drive differentiation of cells into appropriate cell types and structures. This approach has been studied using progenitor cells isolated from both cord and adult blood seeded into the synthetic ECM Matrigel.80 Matrigel constructs implanted into immunodeficient mice for one week demonstrated structures that appeared to be connected to the host vasculature. Although it is not clear that this approach can yield sustained vascular structures, the principle of in vivo directed differentiation and tube formation is supported by these results. A study using an in vivo vascularisation model based on implantation of a chamber around an arteriovenous loop,81 found that Matrigel was able to

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produce vascularised granulation tissue in which the blood vessels did not regress until eight weeks post-implantation. By contrast, synthetic poly(lactic-co-glycolic acid) scaffolds in the same model demonstrated regression of vascularisation by four weeks post-implantation. This is an interesting system as it is able to study the vascularisation process in a semiisolated chamber in the presence of different supporting substrates and active molecules. More recently, the same group showed that infusion of the chemotactic cytokine CXC chemokine ligand 12 enhanced migration of CD34+ endothelial precursor cells into the chamber.82 However, concomitant enhancement of sustained vasculogenesis was not observed since the implantation term was for only seven days. Nevertheless, this approach is a promising one for producing natural vascular structures within a biosynthetic matrix. The challenge of maintaining functional vascularisation remains and it is clear that the presence of appropriate ECM structures is an essential requirement for vascular tube formation. Much of the literature in this field points to the need for ECM components to support and guide formation of vascular structures. However, it is not understood what the most favourable matrix components are for optimal vascularisation. A pertinent model for understanding the development of vascular structures in adults can be found by examining the response to implanted ECM. Livesey et al.83 reported that allogenic acellular dermis implanted into a porcine model was rapidly revascularised. It was proposed that the process used to decellularise the dermis resulted in intact basement membrane (BM) and that other ECM structures were also preserved. The implant was suggested to remain ‘permanently’ and to act as a template for regrowth of dermal structures, including neo-vascularisation. Similarly produced ECM, Alloderm®, is a commercially available decellularised cryo-preserved human dermal matrix. Recent studies of tissue responses to this type of matrix suggest that although revascularisation occurs, the rate is variable depending on the implantation model and there is no clear evidence that existing BM structures associated with native vessels are re-populated.84,85 While there are many approaches being studied for producing vascularised structures in TE constructs, none have successfully produced vascular structures that persist over the longer term. Tissue engineers need to work alongside molecular and developmental biologists to understand the key design criteria for materials that can act as a platform for vascular growth.

7.5

Conclusions

The inability to develop permanent vascularisation in tissue engineered constructs in vivo is a critical barrier that must be overcome before widespread success of TE devices can occur. While natural hydrogels such as

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collagen clearly function as a biological support to developing vasculature, there are issues with fabricating complex devices using such biological polymers. Although many approaches have been used in an attempt to produce robust vascularisation which persists in the long-term, there remains much to learn about how to overcome the natural tendency in adults to form the largely avascular, fibrotic tissue in response to implanted materials. More detailed understanding of embryological development of tissues will also provide critical design parameters for stable vascularisation of tissue engineered devices. Biosynthetic composite hydrogels, with their similarity to natural ECM and the ability to better control their properties, have significant potential for overcoming some of these issues. The challenges for the future in this field revolve around better understanding of the critical ECM and soluble factors for supporting vascularisation, and application of this knowledge to engineer systems that can support specific tissue development.

7.6

Acknowledgements

The authors wish to thank Dr Rylie Green for assistance in preparing the figures, as well as the Australian Research Council (DP0557862 and DP0986447) and the Australia India Strategic Research Fund (DIISR BF010049) for funding.

7.7

References

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8 Inorganic nanoparticles for targeted drug delivery W. PAU L and C. P. S H A R M A, Sree Chitra Tirunal Institute for Medical Sciences and Technology, India

Abstract: Inorganic nanoparticles are non-toxic, hydrophilic, biocompatible and highly stable compared to organic materials. Drug delivery systems designed for enhanced drug efficacy and reduced adverse effects have evolved accompanied by the development of novel materials. Biomedical applications of nanotechnology are mainly suited for diagnostic techniques, nano drugs and delivery systems, and biomedical implants. Nano-enabled drug delivery has been projected as the single largest market opportunity. Recent advancement in nanotechnology has led to the introduction of various inorganic nanoparticles other than calcium phosphates as excellent drug delivery matrices. Nanoparticles are now having highly advanced chemical properties and many inorganic nanoparticles have been used as drug carriers. This chapter reviews some of the recent developments and applications of calcium phosphate nanoparticles, gold nanoparticles and iron oxide nanoparticles in drug delivery and tissue engineering. Key words: inorganic, nanoparticles, drug delivery, calcium phosphate, gold nanoparticles, magnetic nanoparticles, tissue engineering.

8.1

Introduction

Nanostructured materials have unique properties and capabilities that make them suitable for specific interaction with the biological system, particularly in drug delivery applications. Natural systems are made of various nanosized elements that impart unique properties to the biological system. This knowledge has led to investigations and approaches for biomimicking and developing technology for nanomaterials, utilized for different applications. Several organic nanoparticles have been studied and developed successfully from polymers, liposomes and micelles. Drawbacks such as low chemical stability, drug release rate that is unsuitable to the specific application, possibility of microbial contamination, and the undesirable effects of the organic solvents used for particle preparation are some of its inherent problems. Inorganic nanoparticles are non-toxic, hydrophilic, biocompatible and highly stable compared with organic materials. However, less progress has been made in the development of inorganic nanoparticles 204 © Woodhead Publishing Limited, 2010

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in drug delivery systems. Inorganic nanoparticles as drug or gene delivery carriers have received much attention due to their high cellular uptake capacity, non-immunogenic response, and low toxicity. They have exhibited significantly distinct physical, chemical and biological properties from their bulk counterpart’s. Electromagnetic, optical and catalytic properties of noble-metal nanoparticles, such as gold, silver and platinum, are known to be strongly influenced by shape and size. Biomedical applications of metal nanoparticles have been dominated by the use of nanobioconjugates that started in 1971, after the discovery of the colloidal gold labeling technique (immunogold) by Faulk and Taylor (1971). Metal-based nanoconjugates are used in various biomedical applications, such as probes for electron microscopy to visualize cellular components, drug delivery (vehicle for delivering drugs, proteins, peptides, plasmids, DNAs, etc.), detection, diagnosis and therapy (targeted and non-targeted). Drug delivery systems, designed for enhanced drug efficacy and reduced adverse effects, have evolved, accompanied by the development of novel materials. Nanotechnology is an emerging scientific area that has created a variety of intriguing inorganic nanoparticles (Murakami and Tsuchida, 2008; Zhu et al., 2008; Roveri et al., 2008). In recent years, research efforts worldwide have developed nanoproducts aimed at improving health care and advancing medical research. Biomedical applications of nanotechnology are mainly suited for diagnostic techniques, nano drugs and delivery systems, and biomedical implants. Many industries are developing nanotechnology-based applications for anticancer drugs, implanted insulin pumps, and gene therapy. Prostheses and implants are also being developed from nanostructured materials. The global drug delivery products and services market was projected to surpass US$67 billion in 2009. According to the report from NanoMarkets (Nano Drug Delivery, 2005), nanotechnology-enabled drug delivery systems will generate over US$1.7 billion in 2009 and over US$4.8 billion in 2012. Nano-enabled drug delivery is projected as the single largest market opportunity. Reformulations will be possible with the nano-enabled drug delivery systems which help in protecting the patent holders. Recent advancement in nanotechnology has led to the introduction of various other inorganic nanoparticles beside calcium phosphates as excellent drug delivery matrices. They include iron oxide nanoparticles and fullerenes. Carbon nanotubes and nanospheres have been studied as drug delivery vehicles, as their nanometer size enables them to move easily inside the body. The drug can be either inserted in the nanotube or attached to the particle surface. The advantages of inorganic nanoparticles are their very low toxicity profile; biocompatibility and hydrophilic nature; they are not subject to microbial attack and are extremely stable. The importance of these inorganic nanoparticles is ever increasing. It has been established that cells can be taken up by nanoparticles. Thus, nanopar-

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ticles could be utilized to deliver nucleic acids into living cells (Xu et al., 2006). Nanoparticles now have highly advanced chemical properties (Schmid, 1994, 2004; Caruso, 2004; Pileni, 2005) and many inorganic nanoparticles have been used as drug carriers (Bourgeat-Lami, 2002; Chowdhury and Akaike, 2005; Fukumori and Ichikawa, 2006). The inorganic nanoparticles that have been studied in delivering DNA comprise calcium phosphate, carbon nanotubes, silica, gold, iron oxide, quantum dots, strontium phosphate, magnesium phosphate, manganese phosphate, and double hydroxides (Xu et al., 2006; Murakami and Tsuchida, 2008). This chapter reviews some of the recent developments and applications of calcium phosphate nanoparticles, gold nanoparticles and iron oxide nanoparticles in drug delivery and tissue engineering.

8.2

Calcium phosphate nanoparticles

Bioceramics are a class of advanced ceramics that are defined as ceramic products or components employed in medical and dental applications, mainly as implants and replacements. They are biocompatible and can be inert, bioactive and degradable in physiological environments, which makes them ideal biomaterial. Materials that are classified as bioceramics include alumina, zirconia, calcium phosphates, silica-based glasses or glass ceramics and pyrolytic carbons. Calcium phosphates include tricalcium phosphates, hydroxyapatite and tetracalcium phosphates. Calcium phosphate found in bone is in the form of nanometer-sized needle-like crystals of approximately 5–20 nm width by 60 nm length, with a poorly crystallized nonstoichiometric apatite phase containing other trace ions. Unlike tetracalcium phosphates and tricalcium phosphates, hydroxyapatite does not break down under physiological conditions. In fact it is thermodynamically stable at physiological pH and actively takes part in bone bonding, forming strong chemical bonds with surrounding bone. This property has been exploited for rapid bone repair after major trauma or surgery. Since calcium phosphates are biocompatible, resorbable and porous, attempts have been made to utilize them as delivery systems for drugs, chemicals and biologicals. Bajpai and co-workers initiated studies on ceramic drug delivery in the early 1980s by the introduction of alumino calcium phosphorous oxide (ALCAP) ceramic capsules (Bajpai and Graves, 1980; Bajpai and Benghuzzi, 1988). Low cost, ease of manufacture, and biocompatibility makes ceramic materials good candidates for drug delivery applications (Hnatyszyn et al., 1994; Paul and Sharma, 1999). Since hydroxyapatite is biocompatible and is also used as a matrix for the purification of proteins, these particles could be utilized for protein and peptide drug delivery applications. Synthetic calcium phosphates possess exceptional biocompatibility and bioactivity properties with respect to bone cells and tissues, hence have

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been widely used clinically in the form of powders, granules, dense and porous blocks and various composites. For drug delivery, as well as for tissue engineering applications, the present trend is to develop new formulations of hydroxyapatite with properties closer to those of living bone, such as nanosized and monolithic structures. Nanophase calcium phosphate exhibited enhanced osteoblast functions (Webster et al., 2001; Paul and Sharma, 2007) that are very important from the implant application point of view. Nanosized calcium phosphates have been studied in drug delivery systems such as intestinal delivery of insulin (Paul and Sharma, 2001), or of other drugs such as antibiotics (Kano et al., 1994).

8.2.1 Oral insulin delivery applications Type I diabetes is characterized by the inefficiency of pancreatic beta cells to produce insulin. The common form of insulin therapy is by way of twice daily subcutaneous insulin injection. Various attempts have been made to develop a non-invasive delivery system for insulin, namely via oral, buccal, transdermal delivery routes, etc., with varying levels of success. Calcium phosphates have been approved for human use in several European countries as adjuvant. Zinc is also being used for stabilizing insulin (long acting insulins). Therefore calcium phosphates, zinc phosphates and zinc calcium phosphates seem to be suitable candidates for developing ceramic-based insulin delivery systems. Oral cavity delivery is considered to be the most desirable way of delivering drugs from a patient compliance point-of-view. This will be the case with insulin also, once an oral formulation has been developed. Delivered insulin in the case of microspheres follows the same pathway as naturally produced insulin by the pancreas, i.e. reaches directly to the portal circulation and to the liver, consistent with normal physiology. However, in the case of nanoparticles, the absorption of nanoparticles takes place via the Peyer’s patches region, reaches the lymphatic system, bypasses the first pass metabolism (bypasses liver: degradation of insulin is significantly reduced) and the particles will be degraded and delivers the insulin in the blood stream. The concept of oral delivery of insulin utilizing zinc phosphate nanoparticles seems to be promising since insulin will also be stable along with zinc. BioSante Pharmaceuticals, a US based company, has developed calcium phosphate nanoparticles that have successfully passed the first stage of toxicity studies for administration orally, into muscles, under the skin and into the lungs by inhalation. They have been used as a vaccine adjuvant and for protein delivery. Pre-clinical trials of both BioOral and BioAir indicated sustained delivery of insulin with sustained control of glucose levels. Similar formulations may be used for the delivery of other proteins; for example, human growth hormone orally or to the lungs. Calcium phosphate particles containing insulin were synthesized in the

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Blood glucose (% of initial)

140 120 100 80 60 40

CAPIC oral (100 U/kg) Insulin solution oral (100 U/kg) Control

20 0 0

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8.1 Glycemic effect of a single oral dose of CAPIC in fed-diabetic mice (reprinted from Morçöl et al. 2004 with permission from Elsevier).

presence of PEG, and coated with casein to obtain the calcium phosphatePEG-insulin-casein (CAPIC) oral insulin delivery system. When tested in non-obese diabetic (NOD) mice under fasting or fed conditions, it was observed that the biological activity of insulin was preserved by protecting the insulin from degradation while passing through the acidic environment of the GI tract, and it displayed a prolonged hypoglycemic effect after oral administration (Morçöl et al., 2004). Oral administration of 100 U/kg CAPIC to fed-diabetic mice resulted in about 50% reduction of the initial glucose levels within the first 3 h of the treatment, as shown in Fig. 8.1. Glucose returned to the control levels within 5 h. An identical dose of insulin solution (with no calcium phosphate nanoparticles) had no significant effect on blood glucose levels. Calcium phosphate (CaP) nanoparticles with an average particle size of 47.9 nm (D50) were synthesized and the surface was modified by conjugating it with poly(ethylene glycol) (PEG). These modified nanoparticles had a near zero zeta potential. Protection of insulin from the gastric environment was achieved by coating the nanoparticles with a pH sensitive polymer that would dissolve in the mildly alkaline pH environment of the intestine. The release profiles of coated nanoparticles exhibited negligible release in acidic (gastric) pH, i.e. only 2% for CaP and 6.5% for PEGylated CaP. However, a sustained release of insulin was observed at neutral (intestinal) pH for over 8 h. The conformation of the released insulin, studied using circular dichroism, was unaltered when compared with native insulin. The released insulin was also stable as studied using dynamic light scattering.

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100 nm

8.2 Transmission electron microscopy (TEM) image of CaP nanoparticles.

The immunoreactivity of the released insulin was found to be intact. These results suggest PEGylated calcium phosphate nanoparticles as an excellent carrier system for insulin toward the development of an oral insulin delivery system (Ramachandran et al., 2008). A typical transmission electron micrograph of these nanoparticles is shown in Fig. 8.2. They are needle-shaped crystals with 20 nm width and less than 100 nm length.

8.2.2 Gene delivery applications Gene therapy is the insertion of genes into an individual’s cells and tissues to treat a disease, such as a hereditary disease in which a deleterious mutant allele is replaced with a functional one. The first approved gene therapy procedure was performed on 4-year-old Ashanthi DeSilva on September 14, 1990 (Blaese et al., 1995). Doctors removed white blood cells from the child’s body, let the cells grow in the lab, inserted the missing gene into the cells, and then infused the genetically modified blood cells back into the patient’s bloodstream. All viruses bind to their hosts and introduce their genetic material into the host cell as part of their replication cycle. Therefore, the scientists exploited this property and utilized viruses as vehicles to carry ‘good’ genes into a human cell. Viral vectors have high efficiencies in

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delivering the gene and its expressions. However, the drawbacks, such as immunogenecity, cytotoxicity, restricted targeting, production and distribution difficulties, and high costs, led to the eventual clinical failures with viral vectors. Non-viral vectors pose certain advantages over viral vectors, with simple, large-scale production and low host immunogenicity. Several organic nanoparticles have been studied and developed successfully from cationic polymers, liposomes and micelles. Because organic particles tend to microbial attack and were less stable, inorganic nanoparticles such as calcium phosphate got preference. Calcium phosphate nanoparticles represented a unique class of non-viral vectors, which can serve as efficient and alternative DNA carriers for targeted delivery of genes (Maitra, 2005). It has been demonstrated that surface-modified calcium phosphate nanoparticles can be used in vivo to target genes specifically to the liver (Roy et al., 2003). Attachment of a galactose moiety onto the particle surface has increased the targetability of the particles to the liver. This surface modification makes it possible for site-specific gene delivery. Calcium phosphate nanoparticles functionalized by DNA are taken up by living cells. A clear correlation between the uptake of nanoparticles and the efficiency of transfection was found (Sokolova et al., 2007). Surface-modified calcium phosphate nanoparticles by PEGylation encapsulating p53 plasmid have been used for tumor targeted delivery (Fenske et al., 2002). The in vitro transfection studies revealed that consistent levels of gene expression could be achieved by optimizing the Ca/P ratio and the rate of mixing the calcium and phosphate precursors. The optimized forms of calcium phosphate nanoparticles were approximately 25–50 nm in size (when complexed with pDNA) and were efficient at both binding and condensing the genetic material. The differences in gene expression were not just due to a change in size of the naked calcium phosphate nanoparticles but were rather due to the combined effects of pDNA binding and condensation to the particle, which ultimately dictated the overall size of the pDNA–nanoCaPs complex (Oltona et al., 2007). The benefit of calcium phosphate nanoparticles is insignificant IgE response and importantly they are a natural constituent of the human body. Because of these facts, CaP is well tolerated and absorbed by the body. By virtue of the potency of calcium phosphate as an adjuvant and the relative absence of side effects (He et al., 2000), calcium phosphate formulations have great potential for use in humans. Calcium ions play an important role in endosomal escape, cytosolic stability, and enhanced nuclear uptake of DNA through nuclear pore complexes. Calcium phosphate-mediated gene delivery can become more advantageous compared to other viral and non-viral carriers in the sense that the method is relatively safe (since this is in the GRAS (Generally Recognized as Safe) list of FDA), is cost effective, and has high transfection efficiency.

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8.2.3 Cancer chemotherapy applications Hepatocellular carcinoma is the most common primary malignant tumor of the liver. When the human hepatoma cell line BEL-7402 was cultured and treated with calcium phosphate nanoparticles at various concentrations, it inhibited the growth of hepatoma cells in a dose-dependent manner (Liu et al., 2003). The clathrin-mediated endocytosis was found to be responsible for the uptake of calcium phosphate nanoparticles (Bauer et al., 2008). Since the clathrin endocytic pathway keeps the particles captured in the endosomes for lysosomal digestion, the reported toxic effect of calcium phosphate nanoparticles could be caused only by cell structure damage due to accumulating calcium phosphate-filled endosomes, or the toxic effect of lysosomal degraded calcium phosphate solutes in the cytoplasm. Calcium phosphate nanoparticles have been studied for delivery of the chemotherapy agent, cisplatin. The nanoconjugate was prepared by electrostatic binding of cisplatin to the nano calcium phosphate. The drug released from the nanoconjugate was equally effective as the free drug against the A2780cis cell line (Cheng and Kuhn, 2007). A hydrophobic cell growth inhibitor, ceramide, was successfully delivered in vitro to human vascular smooth muscle cells, via encapsulation in calcium phosphate nanoparticles. Nanoparticles encapsulating Cy3 amidite exhibited a nearly five-fold increase in fluorescence quantum yield when compared to the free dye. Thus calcium phosphate nanoparticles can be utilized for encapsulation of imaging agents and anticancer drugs (Morgan et al., 2008).

8.2.4 Tissue engineering applications Synthetic materials could not be considered as ideal implants as the current average lifetime of an orthopedic implant, such as hip, knee, ankle, etc. is only 15 years. Conventional implants, or those implants constituted with grain size dimensions greater than one micron could not invoke natural cellular responses to regenerate tissue that allows longer periods of successful life time. However, since nanophase materials can mimic the dimensions of constituent components of natural tissues, implants developed from nanophase material can be a successful alternative. Several reports on nanophase materials encourage its use for tissue engineering applications. This has been achieved by the combined effect of its ability to mimic the natural nano dimensions, and also the cell responses encouraging high reactivity, which in turn helps in regenerating tissues. NanOss® bone void filler from Angstrom Medica is considered to be the first nanotechnology medical device to receive clearance by the US Food and Drug Administration (in 2005). According to the company, NanOss® is an innovative structural biomaterial that is highly osteoconductive and

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remodels over time into human bone, with applications in the sports medicine, trauma, spine and general orthopedics markets. It is engineered synthetic bone developed from nano crystalline calcium phosphate and is the first material that duplicates the microstructure, composition and performance of human bone. Utilizing nanotechnology, calcium and phosphate are manipulated at the molecular level and assembled to produce materials with unique structural and functional properties. It is prepared by precipitating nanoparticles of calcium phosphate in aqueous phase, and the resulting white powder is compressed and heated to form a dense, transparent, and nano crystalline material. It is strong and also osteoconductive. Ostim® is an injectable bone matrix in paste form, which received CE marking in 2002. It is composed of synthetic nanoparticulate hydroxyapatite, which is indicated for metaphyseal fractures and cysts, acetabulum reconstruction and periprosthetic fractures during hip prosthesis exchange operations, osteotomies, filling cages in spinal column surgery, combination with autogenous and allogenous spongiosa, filling in defects in children, etc. Cell spreading is an essential function of a cell that has adhered to any surface and precedes the function of cell proliferation. Out of the bone and the ceramic material interactions that take place at the material surface, the interaction of osteoblasts is crucial in determining the tissue response at the biomaterial surface (Hunter et al., 1995). Attachment and spreading of specific bone-forming cells in cell culture has been utilized for predicting the behavior of the calcium phosphate materials in vivo (Meyer et al., 2005). The process of cell interaction on materials is highly dynamic and depends on various parameters influencing the cell responses. It is well known that the size and shape of the cell spreading area, as well as the number, size, shape and distribution of focal adhesion plaques, are decisive for further migratory, proliferative and differentiation behavior of anchoragedependent cells (Bacˇáková et al., 2004). Cells usually do not survive if the attached cells are round and are not spreading. If the cell material contact area is significantly high, the cells tend to skip the proliferation phase and enter sooner the differentiation program. If the adhesion is intermediate, the cells are most active in migration and proliferation. Minerals such as zinc and magnesium are known to aid bone growth, calcification and bone density (Higashi et al., 1993; Tucker et al., 1999). Biphasic calcium phosphate ceramic containing zinc also promotes osteoblastic cell activity in vitro (Sogo, 2004). The attachment and spreading of osteoblast-like cells onto porous ceramic materials made from nanoparticles of zinc phosphate and calcium phosphate containing zinc and magnesium has been evaluated (Paul and Sharma, 2007). Since the ceramic matrices are made from nanoparticles, this mimics the way nature itself lays down minerals. An important objective of bone tissue engineering is to develop improved scaffold materials or arrangements to control osteoblast

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behavior significantly affecting its response. Osteoblastic cells on hydroxyapatite exhibited unique attachment and subsequent behavior in vitro, which may explain why mineralized tissue formation is better on hydroxyapatite. Divalent cations, including Mg2+, are known to be active in cell adhesion mechanisms (Fitton, 1995; Hynes, 1992). This investigation demonstrated that the cells were spreading well on matrix containing an optimum amount of calcium, zinc and magnesium. It seemed that the presence of calcium and magnesium encouraged the spreading and adhesivity of osteoblast cells onto nanostructured calcium phosphate matrices. Cells were attached and spread completely on the ZnCaMgP nanomatrix, and this matrix appeared to be comparable to the control group (hydroxyapatite matrix), making it a promising candidate for bone tissue-engineering. Attempts have been made to culture osteoblasts onto nano hydroxyapatite ceramic matrix. The in vitro cultured bone may show further bone-forming capability after in vivo implantation. This tissue-engineering approach is being tried on patients with skeletal problems (Ohgushi et al., 2003). The calcium phosphate nanoparticles, in the order of 100 nm, with high DNA incorporation efficiency, exhibited sustained release of DNA. The MC3T3-E1 preosteoblast cells exhibited the capacity to form bony tissue in as little as two and a half weeks when mixed with DNA nanoparticles encoding for BMP-2 into the alginate hydrogels and injected subcutaneously in the backs of mice, showing its efficacy in bone regeneration applications (Krebs et al., 2009). Similar osteoblast adhesion between nonfunctionalized nano crystalline hydroxyapatite and cell adhesive peptide lysine–arginine–serine–arginine functionalized conventional hydroxyapatite was observed, demonstrating the importance of nano crystalline particles in bone tissue engineering applications (Nelson et al., 2006).

8.3

Gold nanoparticles

In this emerging world of nanoscience, and nanotechnology with nanoparticles, gold nanoparticles, or colloid gold, the most stable metal nanoparticles, have significant importance. The behaviors of the individual particles, and size-related electronic, magnetic and optical properties, are some of the beneficial properties of gold nanoparticles. Colloid gold was used historically for making ruby glass. The most famous example is the Lycurgus Cup that was manufactured in the 5th to 4th century bc. It is ruby red in transmitted light and green in reflected light, due to the presence of gold colloids. The curative power of colloid gold for various diseases has been reported as early as 1618 by the philosopher and medical doctor Francisci Antonii (Antonii, 1618). Colloid gold has been used since the 1930s in modern medicine as a treatment for rheumatoid arthritis. Historically, gold nanoparticles were prepared by citrate reduction from HAuCl4

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50 nm

8.3 TEM image of gold nanoparticles.

to obtain particles in the range of 20 nm. This procedure is the most popular and is still being used for particle preparation. However, several other procedures are also reported by various researchers (Brown and Smith, 1980; Hyatt and Eaton, 1993; Bradley, 1994; Schmid, 1992), giving particles ranging from 1 nm to 150 nm in size. A typical TEM micrograph of gold nanoparticles prepared by reduction of HAuCl4 with sodium citrate in water is shown in Fig. 8.3. Drug, gene, and protein delivery of gold nanoparticles have been reviewed by several investigators (Huang et al., 2007; Ghosh et al., 2008). In an attempt to deliver drug-loaded nanoparticles through the blood– brain barrier, gold nanoparticles were coated with the human serum albumin. The low surface charge and ability to absorb large amounts of creatine, and the albumin layer, may help these particles to attain their objective (López-Viota et al., 2009). Gold nanoparticles, when coated with poly(gamma-glutamic acid) along with phospholipid and polyethylene glycol, exhibited high stability at different pH values and long blood circulation (t1/2 = 22.1 h) upon intravenous injection into mice (Prencipe et al., 2009). Reproducible and reversible phase transition and aggregation of gold nanorod-elastin-like polypeptide nanoassemblies on exposure to nearinfrared light were observed, which finds application in sensing and drug delivery (Huang et al., 2008a). Gold-dendrimer nanoparticles exhibited high levels of uptake and selective targeting to certain organs without specific targeting moieties placed

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on their surfaces (Balogh et al., 2007). Gold nanoparticles were also synthesized and stabilized by new ‘clicked’ dendrimers of generations zero to two (G(0)–G(2)) containing tri- and tetra-ethyleneglycol ethers (Boisselier et al., 2008). Ligand exchanged gold quantum dots conjugated with cellpenetrating peptides are a new class of photoluminescent probes for nuclear targeting and intracellular imaging (Lin et al., 2008). It has been demonstrated that the cellular uptake of functionalized gold nanoparticles with cationic or neutral surface ligands can be readily determined using laser desorption/ionization mass spectrometry of cell lysates. The surface ligands have ‘mass barcodes’ that allow different nanoparticles to be simultaneously identified and quantified at levels as low as 30 pmol. Subtle changes to gold nanoparticles surface functionalities can lead to measurable changes in cellular uptake propensities (Zhu et al., 2008). It has also been demonstrated that incorporating gold nanoparticles in thermosensitive polymer microgels speeds up the response kinetics of PNIPAm, and hence enhances the sensitivity to external stimuli of PNIPAm. These microgels find potential applications for microfluidic switches or microactuators, photosensors, and various nanomedicine applications in controlled delivery and release (Budhlall et al., 2008). PEGylated gold nanoparticle conjugates, which act as water-soluble and biocompatible ‘cages’, allow highly efficient delivery of a hydrophobic drug for photodynamic therapy. The drug delivery time required for photodynamic therapy has been greatly reduced to less than 2 h, compared to 2 days for the free drug (Cheng et al., 2008).

8.3.1 Cancer chemotherapy applications In addition to the surface chemistry of gold nanoparticles, its physical properties were also being exploited for drug delivery applications. Gold nanoparticles can cause local increase in temperature by irradiating it with light in the range of 800–1200 nm. The potential use of gold nanoparticles in photothermal destruction of tumors has been reported by many investigators. Silica–gold nanoshells consisting of a silica dielectric core surrounded by a gold shell coated with temperature sensitive N-isopropylacrylamideco-acrylamide hydrogel modulated drug delivery profiles for methylene blue, insulin, and lysozyme when irradiated by laser. The drug release is dependent upon the molecular weight of the therapeutic molecule (Bikram et al., 2007). Citrate-stabilized gold nanoparticles with a particle size of 30 nm were coated with anti-EGFR (epidermal growth factor receptor) to target HSC3 cancer cells (human oral squamous cell carcinoma). The use of gold nanoparticles enhanced the efficacy of photothermal therapy by 20 times. Colloidal gold has been safely used to treat rheumatoid arthritis for 50 years, and has recently been found to amplify the efficiency of Raman

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scattering by 14–15 orders of magnitude. Large optical enhancements can be achieved under in vivo conditions for tumor detection in live animals. PEGylated gold nanoparticles were brighter than semiconductor quantum dots with light emission in the near-infrared window. When conjugated to tumor-targeting ligand, single-chain variable fragment antibodies, the conjugated nanoparticles were able to target tumor biomarkers such as epidermal growth factor receptors on human cancer cells and in xenograft tumor models (Qian et al., 2008). Polyethylene glycol–gold nanoparticles have been proposed as drug carriers and diagnostic contrast agents. Nanoparticles coated with thioctic acid-anchored PEG exhibited higher colloidal stability than with monothiol-anchored PEG. Coating with high-molecularweight (5000 Da) PEG was more stable. The 20-nm gold nanoparticles exhibited the lowest uptake by reticuloendothelial cells and the slowest clearance from the body; however, they showed significantly higher tumor uptake and extravasation from the tumor blood vessels and are promising potential drug delivery vehicles (Zhang et al., 2009) for tumor therapy. Selective delivery and activity of Kahalalide F analogues have been reported to be improved by conjugating the peptides to gold nanoparticles, for its possible application in antitumor therapy (Hosta et al., 2009). 5-aminolevulinic acid conjugated gold nanoparticles offered a new modality for selective and efficient destruction of tumor cells, with minimal damage to fibroblasts (Oo et al., 2008). Tissue necrosis factor (TNF) has significant therapeutic potential for killing cancer. However, the amount of TNF that patients require has never been delivered successfully without eliciting negative side effects, such as hypotension and in some cases complete organ failure resulting in death. When TNF is coupled with colloid gold, it has been demonstrated that beneficial amounts of TNF can be delivered safely in animal models. CytImmune Sciences Inc. USA uses colloid gold particles that are typically 25 nanometers in size, which is small enough to pass through holes (approximately 100 nm) in the blood vessels that surround a tumor. In healthy organs, spaces between blood vessels are only 5 nm, so colloid gold particles are able to pass into a tumor but are too large to enter any organs. Once the particle passes into the tumor, the TNF is immediately available for biological activity. From the 200–300 TNF molecules around each particle, one molecule acts as the anchor, attaching to cell-surface TNF receptors in and around the tumor, allowing the other TNF molecules to exert their anticancer action. It has been demonstrated that gold-dentrimer nanoparticles had high levels of targeting and this was dependent on surface charge and size (Balogh et al., 2007). Gold-nanoparticle-encapsulated alginic acid-poly[2-(diethylamino)ethyl methacrylate] monodisperse hybrid nanospheres had not only uniform size, similar surface properties, and good biocompatibility, but also unique optical properties provided by the

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embedded gold nanoparticles. These negatively charged nanospheres were internalized by human colorectal LoVo cancer cells and hence could act as novel optical-contrast reagents in tumor-cell imaging by optical microscopy. Drug-loaded nanospheres exhibited similar tumor cell inhibition compared to the free drug doxorubicin (Guo et al., 2009). Functional gold nanoparticles synthesized by a ligand exchange reaction between triphenyl phosphide-stabilized precursor nanoparticles and mercaptopropionic acid were studied with the anticancer drug daunorubicin and were efficient in marking the cancer cells, which may afford potential application for the early diagnosis of the respective cancers (Song et al., 2008). Substantial enhancement of the antiproliferative effect against K-562 leukemia cells of gold nanoparticles (4–5 nm) bearing 6-mercaptopurine9-beta-d-ribofuranoside was achieved compared to the same drug in typically administered free form. The improvement was attributed to enhanced intracellular transport followed by the subsequent release in lysosomes. Enhanced activity and nanoparticle carriers will make possible the reduction of the overall concentration of the drug, renal clearance, and, thus, side effects. The nanoparticles with mercaptopurine also showed excellent stability over one year without loss of inhibitory activity (Podsiadlo et al., 2008). Improved optical coherence tomography image contrast was achieved when gold nanocages were added to tissue phantoms, along with the selective photothermal destruction of breast cancer cells in vitro (Skrabalak et al., 2007). Gold nanoparticles synthesized by employing ‘Gellan Gum’ displayed greater stability to electrolyte addition and pH changes relative to the traditional citrate and borohydride reduced nanoparticles. Anthracycline ring antibiotic doxorubicin hydrochloride loaded nanoparticles showed enhanced cytotoxic effects on human glioma cell lines LN-18 and LN-229 (Dhar et al., 2008).

8.3.2 Gene delivery applications Gold nanoparticles are capable of delivering large biomolecules, without restricting themselves as carriers of only small molecular drugs. Tunable size and functionality make them a useful scaffold for efficient recognition and delivery of biomolecules. They have shown success in delivery of peptides, proteins, or nucleic acids such as DNA or RNA. Complexes formed by using cationic liposomes and polyethylene glycol-modified liposomes, as well as by using phosphatidylcholine liposomes, are proposed for drug and gene delivery (Kojima et al., 2008). Two different DNA oligonucleotides were loaded on two different gold nanorods via thiol conjugation. Selective releases were induced by selective melting of gold nanorods via ultrafast laser irradiation at the nanorods’ longitudinal surface plasmon resonance peaks. Excitation at one wavelength could

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selectively melt one type of gold nanorod and selectively release one type of DNA strand. This seems potential for multiple-drug delivery strategies (Wijaya et al., 2009). Dimethyldioctadecylammonium bromide (DODAB), cationic lipid, bilayer-coated gold nanoparticle efficiently delivered two types of plasmid DNA into human embryonic kidney cells (HEK 293). The transfection efficiency of the gold nanoparticles was about five times higher than that of DODAB (Li et al., 2008a). Stability of this complex with DNA in the presence of serum dramatically increased after coating DODAB onto the surface of the gold nanoparticles. The cytotoxicity of DODAB was also decreased (Li et al., 2008b). Colloidal gold has also been developed as vector for cellular delivery of catalytic DNA enzymes (DNAzymes) which cleave targeted messenger RNA (Tack et al., 2008). Gold nanoparticles were chemically modified with primary amine groups as intracellular delivery vehicles for therapeutic small interfering RNA (siRNA). The resultant core/shell type polyelectrolyte complexes surrounded by a protective PEG shell layer had a well-dispersed nanostructure with a hydrodynamic diameter of 96.3 ± 25.9 nm. The nanosized polyelectrolyte complexes were efficiently internalized in human prostate carcinoma cells, and thus enhanced intracellular uptake of siRNA. Furthermore, the siRNA/gold complexes significantly inhibited the expression of a target gene within the cells without showing severe cytotoxicity (Lee et al., 2008).

8.4

Iron oxide nanoparticles

Magnetic nanoparticles were initially prepared as contrast agents for magnetic resonance imaging. Because of their response to magnetic fields, these nanoparticles seem to be most promising for targeted drug delivery applications. The main disadvantage of the conventional chemotherapeutic drugs is normally their non-specificity. The normal and healthy cells also will be attacked by the cytotoxic drug, along with the tumor cells. Drug-loaded magnetic nanoparticles can be guided towards the tumor cells by the application of a magnetic field, which can reduce the systemic distribution of the drug, thereby reducing the associated side effects. A schematic of the capturing of the magnetic nanoparticles by a magnetic field is shown in Fig. 8.4 (Pankhurst et al., 2003). This can also reduce the drug dosage required for the therapy and efficiently utilize the drug. Once the drug-loaded nanoparticles are concentrated at the tumor site, the drug can be released by applying a stimuli such as enzymatic activity, pH change or temperature (Alexiou et al., 2000). The use of a magnetic field for targeting cancerous tumors was proposed in the late 1970s (Widder et al., 1978; Senyei et al., 1978; Mosbach and Schroder, 1979). The magnetic nanoparticles are usually iron oxide particles and are often applied together with a suitable coating to improve

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Tissue Magnetic nanoparticles

Blood vessel

Tissue

Magnet

8.4 A hypothetical magnetic drug delivery system shown in crosssection (reprinted from Pankhurst et al., 2003, with permission from the Institute of Physics and IOP Publishing Limited).

their biocompatibility and functionalizability. Iron oxide nanoparticles exhibit remarkable new properties, such as superparamagnetism, high field irreversibility, high saturation field, extra anisotropy contributions or shifted loops after field cooling. These properties are contributed by their finite size and surface effects that dominate the magnetic behavior of individual nanoparticles. These nanoparticles have a broad spectrum of applications covering industrial to medical. The nanoparticles intended for biomedical applications should have superparamagnetic behavior at room temperature (Bangs, 1996; Joubert, 1997; Rye, 1996) and should be stable in physiological media. The stability is mainly achieved by their high charge, which avoids precipitation of the particles and settling due to gravitational forces (Langer, 1990). They should also be non-biodegradable in the system (Gupta et al., 2007). Superparamagnetic nanoparticles have also been synthesized as highly monodisperse particles with hydrodynamic diameter of 1.8 nm by a templated synthesis. This film-assisted synthesis gave highly stable particles by preventing aggregation due to magnetostatic coupling (Sreeram et al., 2009).

8.4.1 Cancer therapy applications Therapeutic application of iron oxide nanoparticles include drug-targeting and hyperthermia, where a high degree of specificity is achieved and is mainly applicable to cancer treatments. These particles allowed microparticle retention with an external magnetic field, thus reducing their clearance from the skeletal joint and releasing dexamethasone acetate for 5 days

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in vivo (Butoescu et al., 2009a). Iron oxide nanoparticles, when conjugated with chlorotoxin via amine-functionalized poly(ethylene glycol) silane, enabled the tumor cell-specific binding. They exhibited substantially enhanced cellular uptake and an invasion inhibition rate of approximately 98% compared to unbound chlorotoxin (Veiseh et al., 2009). In another study, iron oxide nanoparticles with maleimidyl 3-succinimidopropionate ligands were conjugated with paclitaxel molecules. These nanoparticles liberated paclitaxel molecules upon exposure to phosphodiesterase (Hwu et al., 2009). Gelatin-coated iron oxide nanoparticles demonstrated a pH responsive drug release leading to accelerated release of doxorubicin at acidic pH compared to neutral pH with increased drug loading efficiency (Gaihre et al., 2009). Starch-coated iron oxide nanoparticles, conjugated with 5-carboxyl-fluorescein (FAM) labeled AGKGTPSLETTP peptide, loaded with doxorubicin, were specific to human hepatocellular carcinoma cell line and exhibited higher cytostatic effect (Yang, 2009). Nanoparticles embedded in a thermo-sensitive Pluronic F127 matrix on short exposure to a high frequency magnetic field caused rapid heating and volume shrinkage of nanospheres, thereby causing instantaneous release of doxorubicin (Liu et al., 2008). Particles embedded in lipid also behaved similarly, releasing the drug under magnetic control (Hsu and Su, 2008). Doxorubicin loaded superparamagnetic iron oxide nanoparticles demonstrated pH dependent release (Munnier et al., 2008). Oleic acid-coated iron-oxide and Pluronicstabilized nanoparticles were loaded with doxorubicin and paclitaxel with a loading efficiency of 90%. These drugs in combination demonstrated highly synergistic antiproliferative activity in MCF-7 breast cancer cells (Jain et al., 2008a). Similar oleic acid-coated nanoparticles (diameter 5–10 nm) were found highly suitable for application towards image-guided drug delivery. These 200 nm particles showed excellent stability under physiological conditions in the presence of fetal bovine serum for 7 days (Talelli et al., 2009). Magnetic particles could be retained in the joint, using an external magnetic field, and prolong the local release of an anti-inflammatory drug. Superparamagnetic iron oxide nanoparticles and dexamethasone 21-acetate (DXM) were co-encapsulated into biodegradable microparticles. DXM encapsulation efficacy was 90%. The microparticles were retained with an external magnet and faster DXM release was obtained for smaller microparticles (Butoescu et al., 2008). Superparamagnetic iron oxide nanoparticles had an excellent biocompatibility with synoviocytes and were internalized through a phagocytic process. This could represent a suitable magnetically retainable intra-articular drug delivery system for treating joint diseases such as arthritis or osteoarthritis (Butoescu et al., 2009b). Iron oxide magnetic nanoparticles functionalized with cisplatin demonstrated synergism between the effects of cisplatin-targetMAG nanoparticles

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and the application of an electromagnetic field (Babincov et al., 2008). Superparamagnetic iron oxide nanoparticles stabilized by alginate exhibited magnetic targeting with an external magnetic field and did not get detained at the injection site without the magnetic field. These nanoparticles were generally considered to be biocompatible and non-cytotoxic (Ma et al., 2008). Biodistribution, clearance, and biocompatibility of magnetic nanoparticles for in vivo biomedical applications were studied by Jain et al. by injecting them intravenously (Jain et al., 2008b) to ensure their safe clinical use. The greater fraction of the injected iron was localized in the liver and spleen rather than in the brain, heart, kidney, and lung. Iron oxide nanoparticles were found to be biocompatible and did not cause long-term changes in the liver enzyme levels or induce oxidative stress. Thus, iron oxide nanoparticles can be safely used for drug delivery and imaging applications. The precise mechanisms of translocation of iron nanoparticles into targeted tissues and organs from blood circulation, as well as the underlying implications of potential harmful health effects in humans, are unknown. One of the latest studies has indicated that the exposure of iron nanoparticles induces an increase in endothelial cell permeability through ROS oxidative stress-modulated microtubule remodeling. This is a new understanding of the effects of nanoparticles on vascular transport of macromolecules and drugs (Apopa et al., 2009). The ability of superparamagnetic iron oxide to be rapidly taken up and distributed into lymphoid tissues was demonstrated for application towards macrophage-targeted nanoformulations for diagnostic and drug therapy (Beduneau et al., 2009). Targeting of mouse xenograft tumors was performed using superparamagnetic iron oxide as a model nanoparticle system. It was found that the elongated nanomaterial (nanoworm) was superior to that of spherical (nanosphere) material in tumor-targeting efficiency and blood half-life (Park et al., 2009). Incorporation of polyethylene glycol as a spacer improved the targeting efficiency. Gold-coated iron oxide nanoparticle/engineered protein G hybrid systems were successfully employed as multifunctional cargo systems for the targeting, imaging, and manipulation of mitochondria (Lim et al., 2009). Improved specificity, extended particle retention and increased cytotoxicity toward tumor cells were demonstrated by iron oxide nanoparticles conjugated with a chemotherapeutic agent, methotrexate, and a targeting ligand, chlorotoxin, through a poly(ethylene glycol) linker. This possesses potential for applications in cancer diagnosis and treatment (Sun et al., 2008). It has been demonstrated with poly(vinyl alcohol) (PVAlc) coated iron oxide nanoparticles that the functionalized nanoparticles displaying potential cellular uptake by human cancer cells depends both on the presence of amino groups on the coating shell of the nanoparticles and of its ratio to the amount of iron oxide (Petri-Fink et al., 2005). Dendrimer-coated iron

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oxide superparamagnetic iron oxide nanoparticles (SPIONs) exhibited selective targeting to KB cancer cells in vitro. The results were consistent between the uptake distribution quantified by flow cytometry using 6-TAMRA UV-visible fluorescence intensity and the cellular iron content determined using X-ray fluorescence microscopy (Landmark et al., 2008). Superparamagnetic iron oxide nanoparticles were functionalized with anti-cancer drugs 5-fluorouridine or doxorubicin via amino PVAlc. The 5-fluorouridine-SPIONs with an optimized ester linker were taken up by human melanoma cells and proved to be efficient anti-tumor agents, while the doxorubicin-SPIONs linked with a Gly-Phe-Leu-Gly tetrapeptide were cleaved by lysosomal enzymes and exhibited poor uptake by human melanoma cells in culture (Hanessian et al., 2008). Intravenously injected iron oxide nanoparticles with hydrodynamic diameters of up to 100 nm were found accumulated in gliosarcomas by magnetic targeting and were successfully quantified by MR imaging. Hence, these nanoparticles appear to be a promising vehicle for glioma-targeted drug delivery (Chertok et al., 2008; Murakami and Tsuchida, 2008; Yu et al., 2008; Peng et al., 2008). Rhodaminelabeled Pluronic/chitosan nanocapsules encapsulating iron oxide nanoparticles were efficiently internalized by human lung carcinoma cells upon exposure to an external magnetic field (Bae et al., 2008). Aminofunctionalized superparamagnetic iron oxide nanoparticles conjugated with Hepama-1, an excellent human monoclonal antibody directed against liver cancer, could markedly kill SMMC-7721 liver cancer cells and could be very useful for bio-magnetically targeted radiotherapy in liver cancer treatment (Liang et al., 2007). Superparamagnetic iron oxide nanoparticles coated with covalentlybound bifunctional poly(ethylene glycol) (PEG) were conjugated with folic acid (FA). The specificity of the nanoconjugate targeting cancerous cells was demonstrated by comparative intracellular uptake of the nanoconjugate and PEG-/dextran-coated nanoparticles by human adenocarcinoma HeLa cells. Uptake of the nanoconjugate by HeLa cells after 4 h incubation was found to be 12-fold higher than that of PEG- or dextran-coated nanoparticles. A significant negative contrast enhancement was observed with magnetic resonance phantom imaging for HeLa cells over MG-63 cells, when both were cultured with the nanoconjugate (Sun et al., 2006). Uptake of methotrexate-immobilized iron oxide nanoparticles, conjugated via a poly(ethylene glycol) self-assembled monolayer, by glioma cells was considerably higher than that of control nanoparticles. The conjugate was highly cytotoxic to 9L cells and the particles internalized into the cellular cytoplasm, retaining its crystal structure therein for up to 144 h (Kohler et al., 2006). Cells expressing the human folate receptor internalized a higher level of methotrexate conjugated nanoparticles than negative control cells (Kohler et al., 2005).

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8.4.2 Gene delivery applications Superparamagnetic iron oxide nanoparticles (SPIONs) were complexed with T cell specific ligand, the CD3 single chain antibody (scAb(CD3), and were used to condense plasmid DNA into nanoparticles with an ideally small size and low cytotoxicity. There was a 16-fold enhancement in the gene transfection level in HB8521 cells, a rat T lymphocyte line (Chen et al., 2009). Iron oxide nanoparticles were coated with carboxylated cholesterol overlaying a monolayer of phospholipid in which Apo A1, Apo E or synthetic amphoteric alpha-helical polypeptides were adsorbed for targeting HDL, LDL or folate receptors, respectively. This could be utilized for in situ loading of nanoparticles into cells for MRI-monitored cell tracking or gene therapy (Glickson et al., 2009). In the presence of magnetic fields, hexanoyl chloride-modified chitosanstabilized iron oxide nanoparticles, conjugated with viral gene (Ad/LacZ), demonstrated high transduction efficiency. The dramatic enhancement in intracellular trafficking of the adenovirus without genetically-modified vesicles can lead to enhanced nuclear transfer, especially in CAR-cells (Bhattarai et al., 2008). Superparamagnetic iron oxide nanoparticles coupled to insulin prevented endocytosis, whereas uncoated particles were internalized by the fibroblasts due to endocytosis, which resulted in disruption of the cell membrane. The derivatized nanoparticles also showed high affinity for the cell membrane (Gupta et al., 2003). Similarly, transferrin-derivatized particles appeared to localize to the cell membrane without instigating receptormediated endocytosis, and also induce up-regulation in the cells for many genes, particularly in the area of cytoskeleton and cell signaling (Berry et al., 2004). This is also the case with lactoferrin or ceruloplasmin-conjugated nanoparticles (Gupta and Curtis, 2004). Iron oxide nanoparticles derivatized with elastin, dextran or albumin induced alterations in cell behavior and morphology distinct from the underivatized particles, suggesting that dermal fibroblast cell response could be directed via specifically engineered particle surfaces. However, there was no difference observed between particles of different sizes (Berry et al., 2002, 2003). It has been demonstrated that higher numbers of conjugated TAT peptide facilitated the cellular uptake of iron oxide nanoparticles in a nonlinear fashion. Cells labeled with these optimized preparations were readily detectable by MR imaging with 100-fold sensitivity (Zhao et al., 2002).

8.4.3 Tissue engineering applications Three-dimensional multicellular assemblies of controlled geometry were formed using external magnetic force utilizing human endothelial pro-

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genitor cells and mouse macrophages magnetically labeled with anionic citrate-coated iron oxide nanoparticles. The procedure avoided the need for substrate chemical or physical modifications. This finds applications in tissue engineering (Frasca et al., 2009). Similarly, monocrystalline iron oxide nanoparticles (MIONs) as an NMR contrast agent have been studied for non-invasive monitoring of tissue-engineered constructs, for optimizing construct design and assessing therapeutic efficacy (Constantinidis et al., 2009). Non-invasive MR monitoring of tissue-engineered vascular grafts (TEVGs) in vivo, using cells labeled with iron oxide nanoparticles has also been demonstrated. This has been achieved by labeling human aortic smooth muscle cells (hASMCs) with ultrasmall superparamagnetic iron oxide (USPIO) nanoparticles. This finds application in the evaluation of in vivo TEVG performance (Nelson et al., 2008). A method to deliver functional nanoparticles and target antibodies into cells was successfully demonstrated with gold-coated iron oxide nanoparticle/engineered protein G hybrid systems (Lim et al., 2009). Targeting of mouse xenograft tumors was performed using superparamagnetic iron oxide as a model nanoparticle system. It was found that the elongated nanomaterial (nanoworm) was superior to that of spherical (nanosphere) material in tumor targeting efficiency and blood half-life. Incorporation of polyethylene glycol as a spacer improved the targeting efficiency (Park et al., 2009). Temperatureresponsive magnetite/polymer nanoparticles developed from iron oxide nanoparticles and poly(ethyleneimine)-modified poly(ethylene oxide)poly(propylene oxide)-poly(ethylene oxide) (PEO-PPO-PEO) block copolymer, with an approximately 45 nm hydrodynamic diameter, underwent a sharp decrease from 45 to 25 nm, when evaluated at temperatures from 20 to 35°C. Thermo-induced self-assembly of the immobilized block copolymers occurred on the magnetite solid surfaces, which is accompanied by a conformational change from a fully-extended state to a highly-coiled state of the copolymer. Consequently, the copolymer shell could act as a temperature-controlled ‘gate’ for the transit of a guest substance. The uptake and release of both hydrophobic and hydrophilic model drugs were well controlled by switching the transient opening and closing of the polymer shell at different temperatures. A sustained release of about 3 days was achieved in simulated human body conditions. In primary mouse experiments, drug-entrapped magnetic nanoparticles showed good biocompatibility and effective therapy for spinal cord damage. Such intelligent magnetic nanoparticles are attractive candidates for widespread biomedical applications, particularly in controlled drug-targeting delivery (Chen et al., 2007). Derivatization of the nanoparticle surface with insulin-induced alterations in cell behavior that were distinct from the underivatized nanoparticles suggests that cell response can be directed via specifically engineered

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particle surfaces. The uncoated particles were internalized by the fibroblasts due to endocytosis, which resulted in disruption of the cell membrane. In contrast, insulin-coated nanoparticles attached to the cell membrane, most likely to the cell-expressed surface receptors, and were not endocytosed. The presence of insulin on the surface of the nanoparticles caused an apparent increase in cell proliferation and viability (Gupta et al., 2003).

8.4.4 General drug delivery and targeting The high biocompatibility and versatile nature of liposomes have been combined with superparamagnetic iron oxide nanocores. The so-called magnetoliposomes find applications in enzyme immobilization for watersoluble, hydrophobic, and other applications such as MRI, hyperthermia cancer treatment and drug delivery (Soenen et al., 2009a). Another result suggests their general applicability, for cell labeling (Soenen et al., 2009b). Iron oxide nanoparticle surfaces were coated with mannan, which induces receptor-mediated endocytosis. Mannan is a water-soluble polysaccharide having a high content of D-mannose residues, to be recognized by mannose receptors on immunate macrophages. This had excellent stability in ferrofluid, and low cytotoxicity. It exhibited enhanced targeted delivery efficiency to macrophages in vitro and in vivo. These nanoparticles are suggested for their potential utility as macrophage-targeting MRI contrast agents (Yoo et al., 2008; Wang et al., 2008). Iron oxide nanoparticles and (18) F-fluoride were encapsulated by hemagglutinating virus of Japan envelopes (HVJ-Es). HVJ-Es were then injected intravenously in a rat and were imaged dynamically using high-resolution PET. Magnetic force altered the biodistribution of the viral envelope to a target structure, and this could enable region-specific delivery of therapeutic vehicles non-invasively (Flexman et al., 2008). Iron oxide nanoparticles covalently bonded to an anti-protein kinase C (PKC)alpha antibody. This conjugate can serve for cellular PKC localization and the inhibition of its function. The activity of the conjugate was confirmed by recognizing PKCalpha using the Western blot method (Makhluf et al., 2008). Lactose-derivatized galactoseterminal-amino iron oxide nanoparticles have been developed for targeting the cell-surface asialoglycoprotein receptor (ASGPR) expressed by hepatocytes (Huang et al., 2008b). AminoPVAlc-SPIONs were taken up by isolated brain-derived endothelial and microglial cells at a much higher level than the other SPIONs, and no inflammatory activation of these cells was observed. Fluorescent aminoPVAlc-SPIONs derivatized with a fluorescent reporter molecule and confocal microscopy demonstrated intracellular uptake by microglial cells.

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Functionalized aminoPVAlc-SPIONs represent biocompatible potential vector systems for drug delivery to the brain, which may be combined with MRI detection of active lesions in neurodegenerative diseases (Cengelli et al., 2006). Nanometer-sized, dextran-coated iron oxide nanoparticles particles were tethered with N-hydroxysuccinimide-folate and fluorescence isothiocyanate (FITC). Internalization of nanoparticles into targeted KB cells occurred only when nanoparticles were conjugated to folate and when the folate receptors were available and accessible on the cells (Choi et al., 2004).

8.5

Conclusion

Because nanotechnology focuses on the very small, it is uniquely suited to creating systems that can better deliver drugs to tiny areas within the body. Nano-enabled drug delivery also makes it possible for drugs to permeate through cell walls, which is of critical importance to the expected growth of genetic medicine over the next few years. The present review discussed the latest ongoing research and development on inorganic nanoparticles utilized as drug carriers and tissue engineering. The targetability and controlled release has been achieved by physico-chemical modification of these nanoparticles. Calcium phosphate nanoparticles are particularly useful as a matrix for bone tissue engineering in addition to their utility in oral insulin delivery systems and non-viral vectors in gene delivery. Iron oxide nanoparticles are extremely effective in active targeting. Similarly, gold nanoparticles can achieve controlled delivery in response to near IR irradiation. However, modifications of these inorganic nanoparticles are required by conjugating with organic materials in order to be effectively used in respective applications, which make them extremely biocompatible. Thus inorganic nanoparticles require an interdisciplinary approach among different fields, including biopolymers, biopharmaceutics, materials science and tissue engineering to enable their potential applications in drug delivery and tissue engineering.

8.6

Acknowledgements

We are grateful to Prof. K. Mohandas, Director, and Dr G. S. Bhuvaneshwar, Head BMT Wing of SCTIMST, for providing facilities for the completion of this work. We are thankful to the laboratory staff for their assistance. This work was partially supported by the Department of Science & Technology, Govt of India, through the project ‘Facility for nano/microparticle based biomaterials – advanced drug delivery systems’ #8013, under the Drugs & Pharmaceuticals Research Programme.

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morçöl, t, nagappan, p, nerenbaum, l, mitchell, a, bell, sjd (2004) ‘Calcium phosphate–PEG–insulin–casein (CAPIC) particles as oral delivery systems for insulin’, Int J Pharma., 277, 91–97. morgan, tt, muddana, hs, altinogˇlu, ei, rouse, sm, tabakovic´, a, tabouillot, t, russin, tj, shanmugavelandy, ss, butler, pj, eklund, pc, yun, jk, kester, m, adair, jh (2008) ‘Encapsulation of organic molecules in calcium phosphate nanocomposite particles for intracellular imaging and drug delivery’, Nano Lett., 8, 4108–4115. mosbach, k, schröder, u (1979) ‘Preparation and application of magnetic polymers for targeting of drugs’, FEBS Lett., 102, 112–116. munnier, e, cohen-jonathan, s, linassier, c, douziech-eyrolles, l, marchais, h, soucé, m, hervé, k, dubois, p, chourpa, i (2008) ‘Novel method of doxorubicin– SPION reversible association for magnetic drug targeting’, Int J Pharm., 363, 170–176. murakami, t, tsuchida, k (2008) ‘Recent advances in inorganic nanoparticle-based drug delivery systems’, Mini Rev Med Chem., 8, 175–183. nano drug delivery (2005) The Impact of Nanotechnology in Drug Delivery: Global Developments, Market Analysis and Future Prospects, NanoMarkets, LC, USA http://www.nanomarkets.net. nelson, gn, roh, jd, mirensky, tl, wang, y, yi t, tellides, g, pober, js, shkarin, p, shapiro, em, saltzman, wm, papademetris, x, fahmy, tm, breuer, ck (2008) ‘Initial evaluation of the use of USPIO cell labeling and noninvasive MR monitoring of human tissue-engineered vascular grafts in vivo’, FASEB J., 22, 3888–3895. nelson, m, balasundaram, g, webster, tj (2006) ‘Increased osteoblast adhesion on nanoparticulate crystalline hydroxyapatite functionalized with KRSR’, Int J Nanomedicine, 1, 339–349. ohgushi, h, miyake, j, tateishi, t (2003) ‘Mesenchymal stem cells and bioceramics: Strategies to regenerate the skeleton’, Novartis Found Symp., 249, 118–127. oltona, d, lic, j, wilsona, me, rogersa, t, closed, j, huange, l, kumta, pn, sfeirc, c (2007) ‘Nanostructured calcium phosphates (NanoCaPs) for non-viral gene delivery: Influence of the synthesis parameters on transfection efficiency’, Biomaterials, 28, 1267–1279. oo, mk, yang, x, du, h, wang, h (2008) ‘5-aminolevulinic acid-conjugated gold nanoparticles for photodynamic therapy of cancer’, Nanomedicine, 3, 777–786. pankhurst, qa, connolly, j, jones, sk, dobson j (2003) ‘Applications of magnetic nanoparticles in biomedicine’, J Phys D: Appl Phys., 36, R167–R181. park, jh, von maltzahn, g, zhang, l, derfus, am, simberg, d, harris, tj, ruoslahti, e, bhatia, sn, sailor, mj (2009) ‘Systematic surface engineering of magnetic nanoworms for in vivo tumor targeting’, Small, 5, 694–700. paul, w, sharma, cp (1999) ‘Development of porous spherical hydroxyapatite granules: Application towards protein delivery’, J Mater Sci Mater Med., 10, 383–388. paul, w, sharma, cp (2007) ‘Effect of calcium, zinc and magnesium on the attachment and spreading of osteoblast like cells onto ceramic matrices’, J Mater Sci Mater Med., 18, 699–703. paul, w, sharma, cp (2001) ‘Porous hydroxyapatite nanoparticles for intestinal delivery of insulin’, Trends Biomat Artif Org., 14, 37–38.

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peng, xh, qian, x, mao, h, wang, ay, chen, zg, nie, s, shin, dm (2008) ‘Targeted magnetic iron oxide nanoparticles for tumor imaging and therapy’, Int J Nanomedicine, 3, 311–321. petri-fink, a, chastellain, m, juillerat-jeanneret, l, ferrari, a, hofmann, h (2005) ‘Development of functionalized superparamagnetic iron oxide nanoparticles for interaction with human cancer cells’, Biomaterials, 26, 2685–2694. pileni, mp (2005) Nanocrystals Forming Mesoscopic Structures, Wiley-VCH, Weinheim. podsiadlo, p, sinani, va, bahng, jh, kam, nw, lee, j, kotov, na (2008) ‘Gold nanoparticles enhance the anti-leukemia action of a 6-mercaptopurine chemotherapeutic agent’, Langmuir, 24, 568–574. prencipe, g, tabakman, sm, welsher, k, liu, z, goodwin, ap, zhang, l, henry, j, dai, h (2009) ‘PEG branched polymer for functionalization of nanomaterials with ultralong blood circulation’, J Am Chem Soc., 131, 4783–4787. qian, x, peng, xh, ansari, do, yin-goen, q, chen, gz, shin, dm, yang, l, young, an, wang, md, nie, s (2008) ‘In vivo tumor targeting and spectroscopic detection with surface-enhanced Raman nanoparticle tags’, Nat Biotechnol., 26, 83–90. ramachandran, r, paul, w, sharma, cp (2008) ‘Synthesis and characterization of PEGylated calcium phosphate nanoparticles for oral insulin delivery’, J Biomed Mater Res B, 18, 699–703. roveri, n, palazzo, b, iafisco, m (2008) ‘The role of biomimetism in developing nanostructured inorganic matrices for drug delivery’, Exp Opinion Drug Del., 5, 861–877. roy, i, mitra, s, maitra, a, mozumdar, s (2003) ‘Calcium phosphate nanoparticles as novel non-viral vectors for targeted gene delivery’, Int J Pharmaceutics, 250, 25–33. rye, pd (1996) ‘Sweet and sticky: carbohydrate-coated magnetic beads’, BioTechnology, 14, 155–157. schmid, g (1992) ‘Large clusters and colloids’, Chem Rev., 92, 1709–1727. schmid, g (1994) Clusters and Colloids. From Theory to Application, Wiley-VCH, Weinheim. schmid, g (2004) Nanoparticles. From Theory to Application, Wiley-VCH, Weinheim. senyei, a, widder, k, czerlinski, c (1978) ‘Magnetic guidance of drug carrying microspheres’, J Appl Phys., 49, 3578–3583. skrabalak, se, au, l, lu, x, li, x, xia, y (2007) ‘Gold nanocages for cancer detection and treatment’, Nanomedicine, 2, 657–668. soenen, sj, hodenius, m, de cuyper, m (2009a) ‘Magnetoliposomes: Versatile innovative nanocolloids for use in biotechnology and biomedicine’, Nanomedicine, 4, 177–191. soenen, sj, vercauteren, d, braeckmans, k, noppe, w, de smedt, s, de cuyper, m (2009b) ‘Stable long-term intracellular labelling with fluorescently tagged cationic magnetoliposomes’, Chembiochem, 10, 257–267. sogo, y, ito, a, fukasawa, k, sakurai, t, ichinose, n (2004) ‘Zinc containing hydroxyapatite ceramics to promote osteoblastic cell activity’, Mater Sci Technol., 20, 1079–1083. sokolova, v, kovtun, a, heumann, r, epple, m (2007) ‘Tracking the pathway of calcium phosphate/DNA nanoparticles during cell transfection by incorporation

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of red-fluorescing tetramethylrhodamine isothiocyanate–bovine serum albumin into these nanoparticles’, J Biol Inorg Chem., 12, 174–179. song, m, wang, x, li, j, zhang, r, chen, b, fu, d (2008) ‘Effect of surface chemistry modification of functional gold nanoparticles on the drug accumulation of cancer cells’, J Biomed Mater Res A, 86, 942–946. sreeram, kj, nidhin, m, nair, bu (2009) ‘Synthesis of aligned hematite nanoparticles on chitosan–alginate films’, Colloids Surf B Biointerfaces. Doi:10.1016/j. colsurfb.2009.02.015. sun, c, fang, c, stephen, z, veiseh, o, hansen, s, lee, d, ellenbogen, rg, olson, j, zhang, m (2008) ‘Tumor-targeted drug delivery and MRI contrast enhancement by chlorotoxin-conjugated iron oxide nanoparticles’, Nanomedicine, 3, 495–505. sun, c, sze, r, zhang, m (2006) ‘Folic acid-PEG conjugated superparamagnetic nanoparticles for targeted cellular uptake and detection by MRI’, J Biomed Mater Res A, 78, 550–557. tabakman, sm, welsher, k, liu, z, goodwin, ap, zhang, l, henry, j, dai, h (2009) ‘PEG branched polymer for functionalization of nanomaterials with ultralong blood circulation’, J Am Chem Soc., 131, 4783–4787. tack, f, noppe, m, van dijck, a, dekeyzer, n, van der leede, bj, bakker, a, wouters, w, janicot, m, brewster, me (2008) ‘Delivery of a DNAzyme targeting c-myc to HT29 colon carcinoma cells using a gold nanoparticulate approach’, Pharmazie, 63, 221–225. talelli, m, rijcken, cj, lammers, t, seevinck, pr, storm, g, van nostrum, cf, hennink, we (2009) ‘Superparamagnetic iron oxide nanoparticles encapsulated in biodegradable thermosensitive polymeric micelles: Toward a targeted nanomedicine suitable for image-guided drug delivery’, Langmuir, 25, 2060–2067. tucker, kl, hannan, mt, chen, h, cupples, la, wilson pw, kiel dp (1999) ‘Potassium, magnesium, and fruit and vegetable intakes are associated with greater bone mineral density in elderly men and women’, Am J Clin Nutr., 69, 727–736. veiseh, o, gunn, jw, kievit, fm, sun, c, fang, c, lee, js, zhang, m (2009) ‘Inhibition of tumor-cell invasion with chlorotoxin-bound superparamagnetic nanoparticles’, Small, 5, 256–264. wang, az, bagalkot, v, vasilliou, cc, gu, f, alexis, f, zhang, l, shaikh, m, yuet, k, cima, mj, langer, r, kantoff, pw, bander, nh, jon, s, farokhzad, oc (2008) ‘Superparamagnetic iron oxide nanoparticle–aptamer bioconjugates for combined prostate cancer imaging and therapy’, Chem Med Chem., 3, 1311–1315. webster, tj, siegel, rw, bizios, r (2001) ‘Enhanced surface and mechanical properties of nano phase ceramics to achieve orthopaedic/dental implant efficacy’, Key Eng Mater., 192, 321–324. widder, kj, senyei, ae, scarpelli, dg (1978) ‘Magnetic microspheres: A model system for site specific drug delivery in vivo’, Proc Soc Exp Biol Med., 58, 141–146. wijaya, a, schaffer, sb, pallares, ig, hamad-schifferli, k (2009) ‘Selective release of multiple DNA oligonucleotides from gold nanorods’, ACS Nano, 3, 80–86. xu, zp, zeng, qh, lu, gq, yu, ab (2006) ‘Inorganic nanoparticles as carriers for efficient cellular delivery’, Chem Eng Sci., 61, 1027–1040. yang, y, jiang, js, du, b, gan, zf, qian, m, zhang, p (2009) ‘Preparation and properties of a novel drug delivery system with both magnetic and biomolecular targeting’, J Mater Sci Mater Med., 20, 301–307.

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9 Alginate-based drug delivery devices L. G R Ø N DA H L, G. L AW R I E and A. J E J U R I K A R, The University of Queensland, Australia

Abstract: Alginates are ubiquitous in the fields of biomaterials science and drug delivery. These polysaccharides are versatile due to their biocompatibility and structural functionality. This chapter provides a comprehensive overview of the recent advances in the development of new alginate-based drug delivery systems through the exploration of the manipulation of chemical and structural properties. The encapsulation of both small molecular drugs and cells producing active biomolecules is considered. The refinement of the assembly of alginate-based matrices for drug delivery has made significant advances and research is now focusing on tailoring the release profiles of the bioactive species. Key words: alginate, drug delivery, encapsulation, ionic and covalent crosslinking, hydrogel.

9.1

Introduction

The investigation of synthetic polymers in the drug delivery field, such as polyethylene and silicon rubber, dates as far back as the 1960s (Desai et al., 1965; Folkman and Long, 1964). However, if the polymer used to construct the delivery system is inert, the drug-depleted delivery system will not break down in the body and will require surgical removal, including associated risks and costs. A more desirable option is the development of drug delivery systems based on polymers that contain hydrolytically or enzymatically labile bonds (Ranade and Hollinger, 2004). Such polymers are termed ‘biodegradable’ polymers and break down in vivo, producing biocompatible products that can be eliminated via normal physiological processes (Ranade and Hollinger, 2004). There are two main categories of such polymers: natural and synthetic, depending on the source from which they are derived. Natural biodegradable polymers include collagen, gelatin, chitosan, alginate, and starch (Ranade and Hollinger, 2004). Each of these polymers offers different advantages, but alginates offer the greatest versatility in their ability to form ionically crosslinked hydrogels. For this reason, alginates form the basis of multiple commercially relevant applications, including drug delivery, tissue engineering, and food technology (de Boisseson et al., 2004; Mørch et al., 2006). 236 © Woodhead Publishing Limited, 2010

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Alginate biopolymers

Commercial alginate is derived from the extracellular matrix of brown algae, where it acts as structural material. Major algae sources are Laminaria hyperborea, Macrocystic pyrifera, and Ascophyllum nodosum. Some other, less frequently used, sources for derivation of alginates are Laminaria digitata and Laminaria japonica (Smidsrød and Skjåk-Bræk, 1990). Alginates are linear unbranched polymers containing β-(1→4)-linked d-mannuronic acid (M) and α-(1→4)-linked l-guluronic acid (G) residues (Fig. 9.1). They are not random copolymers; instead, they consist of blocks of similar and strictly alternating residues (i.e. MMMMM, GGGGG and GMGMGM) (Skaugrud et al., 1999). The distribution of the M and G blocks is dependent on the type and source of the algae, and also on the age and the component of the algae (such as stem or leaves). For example, alginates derived from L. hyperborea have displayed the highest G content, while those from L. japonica are characterized by a low G content; and bacterial alginates produced from the Pseudomonas species are characterized by absence of G blocks. The functional properties of alginates are strongly correlated with the M/G ratio. IR and 1H NMR spectroscopy techniques can be used to determine the M/G ratio in an alginate (Salomonsen et al., 2008). In the IR spectra of alginate, the bands at 1290 and 1320 cm−1 correlate with the M and G content, respectively. Alginate biopolymers may be formulated with a wide range of molecular weights (50–100 000 residues) to suit the application. Enzymes may be used to convert M blocks into G blocks by epimerization, to achieve more control over alginate properties. The discovery of the Azobacter vinelandii, which encodes at least seven different mannuronan C-5 epimerases, opened possibilities of overcoming limitations in applications where high G content was desired (Ertersvag et al., 1994; Svanem et al., 1999). The enzymes use different mechanisms or patterns for conversion of the M blocks into G; for example, AlgE1 enzymes produce long

(a)

(b) CHO

CHO

HO

H

HO

H

H

OH

H

OH COOH

HO

O

6

OH

4 HO

HO

5 3

2

O

HO

H

HO

H

H 1 OH

HO

OH 5 OH

HO

OH H COOH

6 O

4 OH

1

O 3

OH 2

9.1 Chemical structures of the two isomeric monomers of alginate: (a) D-mannuronic acid and (b) L-guluronic acid.

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stretches of G blocks whereas the action of AlgE4 results in strictly alternating sequences. Using the epimerized alginate, Donati et al. made an important discovery of the gelation ability regarding alternating MG blocks with Ca2+ ions, which was previously attributed to G-rich stretches of alginate (Donati et al., 2005). Some other sources of the epimerases have also been identified in the last decade.

9.2.1 Biocompatibility For the purpose of implantation, the biocompatibility of alginate is paramount and has been studied extensively in the literature. Raw alginates contain contaminants such as heavy metals, polyphenols, endotoxins and pyrogens, and immunogenic materials such as proteins that can induce toxic immunogenic responses in the body. In addition, some contaminants are introduced during industrial extraction of alginate or during the gelation reactions, e.g. the chemical compounds used to produce the gels (Nunamaker et al., 2007). Raw alginates stimulate monocytes to produce high levels of cytokines, including interleukin-1, interleukin-6 and TNF-α (Soon-Shiong et al., 1991). It has been reported that the structure of the alginate has an influence on the immunogenic reaction to the alginate implant, with M units of alginate initiating a response in the host in some cases. In separate studies, severe fibrotic overgrowth has been associated with a high G unit content. It was also found that high mannuronic-content alginates were associated with a higher content of polyphenols, endotoxins and proteins. Commercially available alginates can be purified by free-flow electrophoresis to remove these contaminants where they can be separated from the alginic acids under an applied electrical field. Zimmermann et al. (1992) demonstrated that capsules produced from alginates purified by free-flow electrophoresis with high and low M content did not show fibrotic overgrowth when implanted intraperitoneally in rats and mice. Klock et al. (1994) presented an alginate purification method that involved elution and extraction of contaminants from barium crosslinked beads. This method proved very useful for purification of alginate on a large scale. The latter method also has advantages over the free-flow electrophoresis purification in that it does not require the use of sophisticated equipment and reduces the time required for purification. A third method for purifying alginate is based on a multiple step procedure, including the filtration and precipitation of alginate as alginic acid and subsequent extraction of proteins using organic solvents (de Vos et al., 1997). The alginate-PLL microcapsules produced in that study, using the purified alginate, showed improved biocompatibility when implanted in the peritoneal cavity of normoglycaemic rats.

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9.2.2 Degradation Degradation of alginate occurs as a result of cleavage of the glycosidic bonds. This can be influenced or initiated by acidic or alkaline catalyzed reactions, thermal depolymerization or by modification of the polymer such as by oxidation reactions. Alginate is observed to degrade at very acidic and very alkaline pH, whereas it displays high stability at neutral pH. Indeed, it has been proposed by several research groups (Balakrishnan and Jayakrishnan, 2005; Bouhadir et al., 2001; Al-Shamkhani and Duncan, 1995) that alginate is not degradable in vivo unless exposed to a very acidic environment (e.g. gastrointestinal). The alginate polymer does undergo partial hydrolysis at pH 3 in presence of acids such as hydrochloric acid (pKa of M units is 3.38 and pKa of G is 3.65) and depolymerizes at a very alkaline pH, over 10. The M-rich alginates are more susceptible to acid and thermal degradation than the G-rich alginates (Holme et al., 2003). Thermal degradation of the polymer occurs when the temperature is raised above 80 °C. Alginates are also susceptible to degradation in the presence of a group of enzymes called alginases present in marine algae, marine mollusks and microorganisms (Gacesa, 1992). Alginases function efficiently at neutral pH and have been employed in past studies to determine block length and diad frequencies. However, more efficient and sophisticated methods, such as NMR, are employed nowadays for structural characterization of alginate (Section 9.2). Degradation of the polymer can be monitored by techniques such as size exclusion chromatography (SEC) or multiple angle laser light scattering (MALLS), which quantify the molecular weight of the polymer.

9.2.3 Ionically crosslinked alginate hydrogels Ionically crosslinked alginate hydrogels can be formed in the presence of divalent cations such as Ca2+, Ba2+, Sr2+, Zn2+, Cu2+, Cd2+, Co2+ (Mg2+ is an exception) and trivalent cations such as Fe3+ and Al3+ (Remuñán-López and Bodmeier, 1997; Hermes and Narayani, 2002). These gels shrink during gel formation, leading to loss of water and an increase in the polymer concentration relative to the alginate solution. In a study of the correlation between mechanical gel strength and affinity for the cations, Smidsrød and SkjåkBræk (1990) showed that the rigidity of alginate gels in general increased with the affinity of cations, e.g. Pb2+ > Cu2+ > Cd2+ > Ba2+ > Sr2+ > Ca2+ > Co2+ = Ni2+ = Zn2+ > Mn2+. The association between the alginate polymer chain and divalent cations has been popularly described using the ‘egg-box’ model (Fig. 9.2). However, the details of this model have been questioned by many researchers (Braccini and Pérez, 2001; Kim and Han, 2000; Grant et al., 1973;

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Biointegration of medical implant materials Polymer strand 1 O COO HO O

HO

HO

HO

O

O COO− HO O

H

−

HO

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H

H

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O H

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Ca2+

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HO

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O OH O−OC H

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Ca2+ O− OH O OH − O OC O

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Polymer strand 2

9.2 Schematic representation of the interstrand crosslinking of alginate by calcium ions in the so-called ‘egg-box’ model.

Sikorski et al., 2007). The binding involves coordination of one divalent cation with four oxygen atoms. Braccini and Pérez (2001) proposed that the parallel and antiparallel arrangement of 21 helical chains provides a favorable association, also providing a compact cavity for a highly cooperative binding with the divalent cation. More recently, Li et al. proposed 31 helical conformation based on x-ray diffraction of Ca-alginate gels that were formed by a slow gelation method (Li et al., 2007). Divalent cations such as Ca2+ and Ba2+ bind preferentially to the G blocks and thus alginate gels with lower G content (higher M content) have lower strength and stability (Mørch et al., 2008; Jørgensen et al., 2007). Grant et al. (1973) demonstrated that although the interactions of alginate with cations are dominated by G blocks, once the G binding sites are saturated the threshold is passed over to the MG blocks. Thus, a more robust hydrogel can be engineered by using a G-rich alginate or by increasing the amount of G blocks by enzymatic epimerization of the polymer (Mørch et al., 2007). Ionically crosslinked alginate microcapsules can be prepared by various methods (Fig. 9.3) including suspending the extruded alginate beads in a crosslinker solution (Skaugrud et al., 1999; Aslani and Kennedy, 1996) or emulsification followed by addition of a crosslinking solution (Fundueanu et al., 1999). Modifications of the emulsification technique to achieve internal or external gelation have also been reported (Poncelet et al., 1992; Chan et al., 2002). A combination of two methods can also be used to produce desirable gels or to achieve better control of the gelation process (Tan and Takeuchi, 2007). Ionic crosslinking of alginate membranes is commonly

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Sodium alginate solution

Immersion Incorporated calcium carbonate

Calcium chloride solution

Acidic solution

Injecting barium chloride crystals

Barium chloride solution

9.3 Illustration of the common techniques used to produce ionically crosslinked alginate capsules.

achieved by the immersion technique, where the xerogel is placed in a crosslinker solution allowing the crosslinking agent to diffuse into the matrix (Corkhill et al., 1990; Pavlath et al., 1999; al Musa et al., 1999; Jejurikar et al., 2008). Recently a new method of pressure-assisted diffusion was developed to achieve more uniform crosslinking (Jejurikar et al., 2008). The internal structure of alginate hydrogels can be investigated by cryogenic SEM, which allows evaluation of pore size and distribution of the hydrated gel (Fig. 9.4). These ionically crosslinked hydrogels swell significantly in an aqueous media. Ca-alginate capsules are known to swell more than 90% over long-term immersion in physiological media under optimal pH and temperature conditions. The extent of swelling of a crosslinked alginate gel is dependent on temperature, pH, and ion concentration of the swelling media, and also on the crosslinking gradient through the gel matrix (Moe et al., 1993). In a recent study, Qin reported that Ca-alginate gels with higher M content swell more than those with higher G content, which reflects the preferential binding of the cations to the G sub-units (Qin, 2008). The mechanical properties of alginate hydrogels change over an initial period of immersion in a physiologically relevant solution (Le Roux et al., 1999; Kuo and Ma, 2008) and are dependent on the technique and material used to prepare the hydrogels, as illustrated in Table 9.1. The stress–strain response of the hydrogel is dependent upon the extent of crosslinking, the porous structure of the matrix and the water flow within.

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9.4 Cryogenic scanning electron micrograph of a hydrated calcium crosslinked alginate matrix showing its porous nature.

Table 9.1 Mechanical properties of calcium crosslinked alginate hydrogels Alginate sample and method for preparation Capsules prepared by extrusion method (Nguyen et al., 2009) Cylinders prepared by internal gelation method (le Roux et al., 1999) Discs prepared by emulsion method (Webber and Shull, 2004) Discs prepared by internal gelation method (Drury et al., 2004) Films prepared by immersion (Jejurikar et al., 2008) Films prepared by pressure assisted diffusion (Jejurikar et al., 2008)

Mechanical properties

Maximum modulus (kPa)

Compression

490

Compression

10

Compression

100

Tensile

50

Tensile

5 400

Tensile

16 500

Degradation of ionically crosslinked alginate gels is initiated by loss of crosslinking cation, followed by loss of high and low molecular weight polymer strands. This affects the gel strength, stability, pore size and pore distribution (Shoichet et al., 1995). The rate of degradation is affected by the alginate properties (in particular the M/G ratio) and also by the solution

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in which the alginate gel is suspended (Tan et al., 2009). The effect of the calcium concentration in the medium was investigated in detail by Kuo and Ma (2008) and it was found that a high concentration (5 mM) resulted in greater retention of the crosslinking density than a lower concentration (e.g. 2–4 mM). It is of course important to realize that the mechanism and rate of degradation in vivo is also dependent on the site of implantation and the local cellular environment.

9.2.4 Chemically modified alginate Chemical modification of the alginate biopolymer can change the reactivity of alginate either by introducing highly reactive functional groups (e.g. aldehyde groups) or by introducing chemical (e.g. phosphate) or biochemical (e.g. amino acids) groups that can increase the biointegration of alginate-based materials. Since the alginate biopolymer itself contains both hydroxyl and carboxylic acid functional groups, it offers great versatility for chemical modification, as will be illustrated below. It has recently been demonstrated that phosphorylation of alginate can be achieved using the so-called urea phosphate method (Coleman, 2007). This entails reacting a dispersion of the biopolymer with phosphoric acid in DMF in the presence of urea (Mucalo et al., 1995). In this manner, phosphate groups are introduced at C1, C2 and C3 through reactions with the hydroxyl groups. Characterization by NMR (1H, 13C and 31P) established that the dominant site of phosphorylation on d-mannuronic acid residues was M-3, which indeed is the equatorial group and therefore more reactive than its axial counterparts due to reduced steric hindrance. Because of the high acidity of the solution during the phosphorylation reaction, some degradation of the biopolymer occurs concurrently with the phosphorylation reaction. It was thus possible to achieve up to 26% phosphorylation (i.e. 26% of all alginate subunits carry a phosphate group) and this is paralleled with a decrease in molecular weight from approximately 140 to 40 kDa. Coupling of single amino acids (e.g. cysteine) or small peptides (e.g. containing the RGD sequence) to alginates has been studied in detail. In all cases the amine terminal of the amino acid/peptide is reacted with the carboxylic acid group of the alginate using carbodiimide chemistry (i.e. activation of the carboxylic acid group). More specifically, cysteine, has been coupled to alginate with 7% substitution (of alginate subunits, 400 μmol/g) when using stoichiometric amounts of cysteine, as determined photometrically (Greimel et al., 2007). The aim was to improve the mucoadhesive properties of ionically crosslinked capsules used for oral delivery (Section 9.3.2). Other amino acids including lysine, arginine, aspartic acid and phenylalanine have likewise been coupled to alginate (Zhu, 2002). Successful reaction was verified from FTIR and 1H NMR; and the amino acid content

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was found to be 3% (alginate subunits) based on analysis by UV spectrometry. The RGD peptide sequence is used extensively in biomaterials science due to its cell adhesion properties. A number of RGD-containing peptides have been coupled to alginate using carbodiimide chemistry; including GRGDY (Rowley et al., 1999), GGGGRGDSP (Drury et al., 2005), and GGGGRGDSY (Connelly et al., 2007). The coupling efficiency was determined using 125I labeled peptide and found to be 25% (alginate subunits) without further improvement with increased carboxylic acid activation (Rowley et al., 1999). These peptide-modified alginate biopolymers have found many applications, some of which will be discussed in the following sections of this chapter. Amphiphilic alginate derivatives have been synthesized by reacting the carboxylate groups of alginate, transformed into a tertrabutyl ammonium salt, in the organic solvent dimethylsulfoxide with alkyl halides (e.g. dodecyl bromide) (Pelletier et al., 2000). In this manner, the long alkyl chains were linked to the alginate biopolymer via ester linkages. An 8% (C12 alkyl chain length) and 1.3% (C18 alkyl chain length) substitution ratio could be achieved in this manner, as determined through gas chromatography of the alkyl alcohol obtained from alkaline hydrolysis of the amphiphilic polymer. An alternative method of reacting the carboxylic acid group of alginate, pre-activated using chloro-1-methylpyridinium iodide (CMPI), with an alkylamine (e.g. dodecylamine) in dimethyl formamide yielded substitution ratios in the range of 2–17% and this could be easily controlled by the amount of CMPI used (Vallee et al., 2009). Such amphiphilic alginate biopolymers can form physical gels due to hydrophobic interactions, and these gels can be further strengthened by calcium crosslinking (de Boisseson et al., 2004). Oxidation of alginate using periodate was first described by Malaprade (1928). It involves cleavage of the C2–C3 bond, transforming the uronic acid subunits into an open chain adduct containing a dialdehyde known as alginate dialdehyde (ADA) (Fig. 9.5). These aldehyde groups can react spontaneously with hydroxyl groups present on the adjacent uronic acid subunits in the polymer chain to form a hemiacetal (Balakrishnan et al., 2005). Its formation can prevent complete oxidation of the biopolymer and consumption of sodium periodate. The open chain adduct created by the periodate oxidation of alginate is much more susceptible to hydrolytic scission than unmodified alginate (Bouhadir et al., 2001). Oxidation of alginate is usually carried out in an aqueous solution at room temperature (Balakrishnan et al., 2005; Gomez et al., 2007). The degree of oxidation can be controlled, and up to 87% oxidation has been demonstrated (Balakrishnan et al., 2005). The degree of oxidation can be determined via iodometry by titrating any periodate remaining after reaction (Whistler and Wolfrom, 1962). Characterization is carried out using FTIR (formation of the alde-

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Alginate-based drug delivery devices –

–

OOC

OH

OHO

OH –

OOC O

O

OH

OOC

245

OH O

O HO

Sodium periodate (aqueous) –

–

OOC

OHO

OH

–

OOC O

CH O

O HC

OOC

O HO

OH O

:O:

9.5 Reaction scheme of the oxidation reaction of alginate forming dialdehyde alginate (ADA).

hyde and acetal bands), 1H NMR (aldehyde content) (Gomez et al., 2007), and 13C NMR (to confirm the polymer scissioning that is accompanied by oxidation). Extensive scissioning of the polymer chain is reflected in the decrease in molecular weight with an increase in the amount of oxidation. Vold et al. (2006) and Smidsrød and Painter (1973) demonstrated changes in the intrinsic viscosity and chain stiffness in the oxidized alginate samples and concluded that the stiffness and the length of chains were decreased with an increase in the oxidation of alginate samples. The enhanced flexibility of the polymer chains was linked to the open-chain adducts. The gelation ability of ADA in the presence of Ca2+ ions is maintained at low oxidation levels (e.g. 10%); however, the Ca-ADA hydrogels show reduced mechanical properties. This is related to the reduced molar mass and number of GG blocks participating in binding with divalent cations. ADAs are generally biocompatible; however, alginates oxidized to a very high degree can induce an immunological response as a consequence of the presence of free radicals. It has been reported that alginates are resistant to biodegradation and are not broken down in mammals (Al-Shamkhani and Duncan, 1995). However, a controlled degradation of alginate can be achieved by partially oxidizing the alginate polymer, which opens avenues for use of ADA in tissue engineering applications (Bouhadir et al., 2001).

9.2.5 Covalently crosslinked matrices Covalently crosslinked hydrogels, possessing increased stability compared to their ionically crosslinked counterparts, allow for a greater control over mechanical and swelling properties of the gels, and can be engineered to suit various applications such as cartilage repair, wound dressing and drug

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delivery. As discussed above, the hydroxyl and carboxylic acid functional groups of alginate offer great versatility for chemical modification, including crosslinking reactions. Formation of covalently crosslinked hydrogels can thus be achieved by reaction of these functional groups with complementary reactive groups (Augst et al., 2006). Crosslinking molecules such as dialdehydes and diamines have been used with suitable catalysts to create hydrogels. A crosslinking reaction between the hydroxyl groups of alginate and the aldehyde groups of typically gluteraldehyde (1,5-pentanedial) is one approach that has been studied extensively (Chan et al., 2008; Yeom and Lee, 1998). The chemical reaction leads to the formation of acetal groups in the crosslinked network. Kim et al. demonstrated that gluteraldehydecrosslinked alginate hydrogel fibers exhibit high absorbency (Kim et al., 2000), which could be increased by decreasing the crosslinker concentration to just above the critical concentration required for crosslinking of the polymer in agreement with general hydrogel theory. These hydrogels have potential application in products such as additives for soil in agriculture, water-blocking tapes, sanitary napkins, disposable diapers or drug delivery systems, where water retention is important, or as membrane material for pervaporation processes. However, gluteraldehyde is toxic to cells even in low concentration and hence these hydrogels are not ideal candidates for biomedical applications. Alternatively, alginate dialdehyde (ADA) obtained by periodate oxidation of alginate can be used as the aldehyde crosslinker to react both intra-molecularily with other parts of the ADA and intermolecularily with unmodified alginate (Jejurikar et al., 2009). These gels displayed very high swelling characteristics and good mechanical strength which could be controlled by the degree of oxidation of ADA, as well as the ratio of ADA to alginate. ADA has also been crosslinked with adipic acid hydrazide. The swelling and degradation properties of these hydrogels were controlled by varying the amount of adipic acid hydrazide in the reaction mixture (Lee et al., 2004). Direct crosslinking of hydroxyl and carboxylic acid groups of alginate using the non-toxic carbodiimide (EDC) catalyst has been demonstrated (Rowley et al., 1999; Xu et al., 2003a). EDC facilitates formation of ester linkages through an intermediate and does not itself remain a part of the structure. Xu et al. (2003b) made blends of alginate and carrageenan and used EDC chemistry to create more durable hydrogels. These hydrogels predominantly contained alginate, thus preserving their water absorbing properties. Xu et al. demonstrated that these hydrogels paralleled the gluteraldehyde-crosslinked hydrogels as good candidates for membrane materials for the pervaporation process (Xu et al., 2003a). EDC chemistry, when coupled with the co-reactant N-hydroxyl succinimide (NHS), has been utilized to facilitate amide bond formation between the carboxylic acid groups

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of alginate and the amine groups of diamines. This strategy has been utilized with methyl ester l-lysine, ethylene diamine, and poly(ethylene glycol) (PEG) diamines (Lee et al., 2000a; Hosoya et al., 2004; Eiselt et al., 1998, 1999). Eiselt et al. developed alginate hydrogels with demonstrated improved mechanical properties, compared to ionically crosslinked alginate hydrogels, by crosslinking alginate and PEG-diamines (Eiselt et al., 1998, 1999). The mechanical properties were controlled by the molecular weight of the PEG and the amount of PEG-diamines introduced into the crosslinked network. Alginate crosslinked with ethylenediamine and/or an amine silane have been reported as useful for making bone-like apatite, artificial skin, and nerves (Hosoya et al., 2004). A number of other methods have also been explored for the covalent crosslinking of alginate, including irradiation and enzymatic techniques. Alginate has been modified using methacrylic anhydride (Smeds and Grinstaff, 2001), leading to an ester linkage between alginate and the vinyl-group containing moiety. These modified polymers in the presence of a photoinitiator and UV irradiation produced covalently crosslinked hydrogels exhibiting desirable swelling and mechanical properties. These gels are of great interest for sutureless closure of wounds and tissue reconstruction/ artificial organs. A recent example of enzymatically crosslinked alginate utilizes tyramine-modified alginate (coupled via amide linkages using EDC/ NHS chemistry) (Sakai and Kawakami, 2008). Horseradish peroxidase was applied to catalyze the oxidation reaction, coupling the phenols, thereby creating covalent crosslinks. This was done on calcium crosslinked gels and it was found that the swelling was reduced and could be controlled by the reaction conditions.

9.3

Drug delivery using alginate matrices

The transport of pharmaceuticals to the desired target site and their release in a controlled and sustained way represents a massive research field of drug delivery. To simply focus on alginate-based systems and claim that these represent the state of the art in this field would be naïve. However, a number of strategies in drug delivery can be illustrated through important examples in polysaccharide-based systems. Drug delivery capsules are designed with two key objectives in mind: the safe transport of the active molecule to the active target site and minimal impact on the host biological system. Typically, engineering of the structure of the assembly is either directed towards improved permeability, stability and retention (passive targeting) or the modification of the functionality of the capsule to improve localized delivery and/or biointegration (active targeting). A number of these strategies are illustrated in Fig. 9.6 and will be discussed further below (Sections 9.3.1 and 9.3.3).

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Biointegration of medical implant materials Key Alginate

Modified alginate Calcium ion

Cationic polyelectrolyte Anionic polyelectrolyte Covalently crosslinked polyelectrolyte complex

(e) LbL coating

(a) (d)

(f)

(b)

(c)

9.6 Schematic illustration of various approaches used to control permeability and biointegration. (a) Simple crosslinked alginate matrix; (b) crosslinked matrix produced from chemically modified alginate; (c) post modification of alginate matrix with chemical or biological moieties; (d) LbL assembly around alginate matrix; (e) covalent crosslinking of LbL assembly; (f) post modification of LbL assembly.

The strategies that are implemented depend on the nature of the drug to be delivered and the target environment (Table 9.2). Oral delivery of drugs represents a particular challenge due to the complex gastrointestinal environment, and there are several examples of research directed towards acquiring the desired characteristics for efficacy via this delivery route. These strategies include: modification of alginate to introduce binding through disulfide bonds with mucous glycoproteins (Greimel et al., 2007); novel crosslinking strategies to strengthen capsules (Anal and Stevens, 2005); application of a covalently crosslinked chitosan coat (Taqieddin and Amiji, 2004), and introduction of buoyancy to evade gastric emptying (Shishu et al., 2007). The layer-by-layer (LbL) assembly approach has been utilized for a number of different applications with the aim of modulating the release profile of the encapsulated drug through the formation of a diffusional barrier around the alginate capsule. In the work by Matsusaki

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Table 9.2 Common drug delivery environments Delivery site

Desirable properties

Gastrointestinal

Stability in low pH environment. Non-covalent associations with mucous gel layer. Prolonged gastric residence time. Enhanced lifetime within circulatory system. Small size to pass through capillaries. High surface area to volume ratio to maximize contact. Efficient transfer to blood or lymphatic system. High volume and sustained release. Bioresorbable or biomimetic.

Intravenous Peritoneal cavity Subcutaneous Hard tissue fracture

et al. (2007) encapsulating vascular endothelial growth factor (VEGF) in alginate capsules, it was found that varying the number of layers in the LbL assembly had a significant effect on the rate of VEGF release. Modification of microcapsules to reduce the local inflammatory or tissue response in the host represents an emerging area of interest in biointegration (Section 9.3.3). Strategies include the modification of alginate to incorporate antibodies on the outer surface of the microcapsule, e.g. anti-TNF-α (Leung et al., 2008) and heparin as the outer layer of a multilayered capsule assembled by physical adsorption (Bünger et al., 2003).

9.3.1 Use of chemically modified alginate The uses of chemically modified alginate in drug delivery applications have been explored for different aspects of biointegration. The nature of the physicochemical relationship between drug molecule and carrier matrix impacts on the drug release profile and so the design of capsules represents a balance between the matrix/environment and drug/matrix associations. Examples considered here include the use of thiolated alginate for mucoadhesion, the use of amphiphilic alginate derivatives for enhancing protein retention, and the use of oxidized alginate for both formation of a covalently crosslinked matrix and covalent attachment of a drug molecule. While the first example is mainly driven by the interactions of the delivery system with the in vivo environment, the last two examples explore the drug/matrix interactions. Thiomers have been proposed to be the next generation of mucoadhesive polymers to be used in oral drug delivery applications (Bernkop-Schnürch and Greimel, 2005). They have been proposed to act by thiol/disulfide bond exchange reactions and oxidation reactions with cysteine-rich subdomains of mucus glycoproteins, thus forming covalent linkages to mucin via

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disulfide bonds. In their study on producing mucoadhesive alginate delivery capsules, Greimel et al. (2007) used mixtures of thiolated alginate (modified by cysteine attachment, Section 9.2.4) and poly acrylic acid (PAA) to encapsulate insulin. They found that the introduction of the thiol group offered additional benefit to the delivery system. It increased the encapsulation efficiency from 15% (unmodified alginate) to 65% (thiolated alginate/ PAA) and decreased the release rate. In addition, capsule stability in simulated intestinal fluid was significantly enhanced and this was attributed to the formation of internal disulfide bonds during capsule formation (i.e. 82–85% of free thiol groups were oxidized). These particles thus show some promising characteristics for use in oral drug delivery. Hydrophobically modified alginates have been trialled for encapsulation of proteins through the incorporation long-chain alkyl groups, e.g. C12 via ester bonds (Leonard et al., 2004). These amphiphilic alginate derivatives were used to encapsulate a series of model proteins with very high encapsulation efficiencies (70–100%), in which both hydrophobic interactions and calcium ion crosslinking were contributing to the bead formation process. In addition, the strong interactions between the proteins and the amphiphilic alginate derivatives were illustrated by retention of the proteins in simple buffers. This was in contrast to the rapid release of these proteins from unmodified alginate capsules. Protein release from the amphiphilic capsules was observed only in the presence of the esterase lipase, which hydrolyzes the ester bond, or in the presence of surfactants that disrupt the hydrophobic interactions in the capsule. An elegant use of oxidized alginate (alginate dialdehyde, ADA) in drug delivery has been demonstrated by Balakrishnan and Jayakrishnan (2005). They produced an injectable scaffold from ADA and gelatin in the presence of small amounts of borax. It was observed that the gelling time decreased with increasing concentration of all reagents, while it increased with the degree of oxidation of ADA. Using primaquine as a model drug, they demonstrated slower drug release rates when using ADA of a high degree of oxidation and attributed this to the ability of the drug to undergo a Schiff’s reaction with the aldehyde groups, requiring subsequent degradation (rather than simple diffusion) processes to take place in order to release the drug.

9.3.2 Composite matrices for targeted delivery Biocomposites are widely used in bone repair and regeneration applications (Grøndahl and Jack, 2009). Due to their osteoconductive properties, hydroxyapatite (HAP) and other calcium phoshate phases have been widely researched for such biocomposite fabrication. A number of studies have shown promising in vitro bioactivity of HAP/alginate composite matri-

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ces for bone tissue engineering applications (Lin and Yeh, 2004; Turco et al., 2009). Maruyama (1995) investigated the in vivo performance of HAP/ alginate composite capsules and found them to show excellent osteoconductivity properties, bridging the gap between the implants and the cortical bone without adverse effects. In addition, they proposed that uniformly packed spherical particles with monodisperse porosity supported the rate of bone growth. Sivakumar and Rao (2003) applied this knowledge to the development of a delivery system for the antibiotic gentamicin, using coralline hydroxyapatite microspheres to encapsulate the drug. With the same general aim of drug delivery to bone fractures, the antimicrobial drug Biocide 1 was encapsulated in HAP/alginate granules (Krylova et al., 2002); the antibiotic gentamycin was encapsulated in HAP/alginate particles (Paul and Sharma, 1997); the therapeutic enzyme glucocerebrosidase was encapsulated in HAP/alginate and calcium titanium phosphate/alginate microspheres (Ribeiro et al., 2004); and a glucosaminoglycan was encapsulated in HAP/alginate capsules (Tan et al., 2009). Among the studies reporting drug delivery from HAP/alginate systems, only two include a comparison to the release from a pure alginate system (Ribeiro et al., 2004; Tan et al., 2009). In the work by Ribeiro et al. (2004), it was reported that the protein drug was released more rapidly from the pure alginate capsules than from composite ones when the protein was pre-adsorbed to the inorganic particles. This was attributed to the high affinity of HAP for the protein. A similar effect was observed in the study encapsulating heparin, where some retention of the drug was observed for the composite capsules (Tan et al., 2009). However, the systems in this study were very complex and the release rate was affected by the different types of intermolecular interactions between the glucosaminoglycan molecules and the alginate and HAP matrix components.

9.3.3 Microencapsulation for transplantation The transplantation of cells is a viable strategy in drug delivery, where therapeutic cells can be derived either from the same species (allograft) or, more commonly, a different species (xenograft). The relatively high biocompatibility and low cytotoxicity of purified alginate have made it a popular biomaterial for the microencapsulation, xenotransplantation and immunoisolation of these cells. The provenance of this field can be traced to the research of Lim and Sun who, in 1980, reported the assembly of a calcium alginate–polylysine–alginate (APA) microcapsule around pancreatic islets (Lim and Sun, 1980). A large body of the subsequent literature has been directed towards attempts to create the bioartificial pancreas through the microencapsulation and immunoisolation of pancreatic islets; there are multiple reviews relating to this specific application (Wilson and Chaikof,

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2008; Lacik, 2006; de Vos et al., 2006; Narang and Mahato, 2006). One of the hurdles to progression in this field has been the diversity of methodologies and materials adopted for microencapsulation, resulting in a consensus that controlled systematic studies are required. In recent years, researchers have collaborated towards this goal, catalyzed by the EU-COST project (http:// cost865.bioencapsulation.net/). Despite the variability in methodologies, the substantial number of studies published prior to 2006 has enabled refinement of a suite of desirable and undesirable properties in alginatebased microcapsules which contribute to their potential success in their application (de Vos et al., 2006; Lacik, 2006; Zimmermann et al., 2005). These characteristics have been summarized in Table 9.3. Despite the demand for the standardization of methodologies surrounding the microencapsulation of pancreatic islets, it is evident that the microcapsule matrix must be individually tailored for each cellular application (Fig. 9.7). One of the challenges in optimizing the microcapsule for a desired application relates to the strength and permeability of the alginate-based matrix, with undesirable outcomes including rupture or the escape of cells during proliferation and differentiation (Li et al., 2008; Lee et al., 2009). The porosity and strength of the matrix can be manipulated through the selection of the type of alginate, with the ratio of M/G units dictating the extent of crosslinking (refer to Section 9.2.3). There are multiple approaches that have been adopted for a wide range of cellular systems and target environments, a summary of the most recent studies applying microencapsulation in alginate-based matrices being provided in Table 9.4 demonstrating the diversity and expansion of this therapeutic approach. While there is advancement in the optimization of the matrices, these studies continue to raise challenges such as promotion of the process of vascularization to optimize oxygen exchange, which may be overshadowed by the detrimental outcome of fibrotic overgrowth (Wilson and Chaikof, 2008). A survey of the strategies summarized in Table 9.4 reveals that there are several levels of complexity in the structural matrix of the microcapsules being trialled. These include: simple crosslinked gels (Fig. 9.6a); multilayered self-assembled polyelectrolyte complexes (Fig. 9.6d); and post modification to introduce desirable attributes (Fig. 9.6e and f). Each of these levels is described in detail below. 1

Alginate hydrogels ionically crosslinked by divalent cations are the simplest systems that have been adopted for cell microencapsulation and, of these, the most widely adopted continues to be Ca-alginate hydrogels. The process of forming hydrogels around the cells by this approach is very gentle, with low shear stresses and minimal exposure to reagents. The characteristic properties of the type of alginate selected for this purpose have a significant impact on the subsequent viability

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Table 9.3 Desirable attributes in the assembly of alginate-based microcapsules Desirable

Undesirable

Alginate of sufficient purity so that it does not induce a cytotoxic response in the hosted cells. Alginates with an endotoxin content >100 EU/g are not suitable for in vivo studies. A robust structure which can withstand localized compression and shear stresses imparted by the target environment. Longevity of the matrix resistant to biodegradation. Tailoring of the capsule through a blend of a high G alginate imparting mechanical strength and a high M alginate promoting elasticity when required. A semi-permeable membrane which can retain the cells that it surrounds and simultaneously permits diffusion of both the nutrients and the active biomolecules. Oxygen transport to the encapsulated cells is a critical process. The pore dimensions and their interconnectivity in the polymer network control diffusion.

Toxins inherent in unpurified alginates include pyrogens, mitogens, polyphenols, and peptides. These substances invoke recognition by macrophages and induce fibrotic overgrowth of the microcapsules. A low degree of cross-linking results in a structure which swells and degrades in the presence of biological fluids. This also relates to the M/G ratio of the alginate and the viscosity. In contrast, a membrane that is too rigid (high degree of cross-linking) ruptures easily.

The outer surface of the microcapsule should be immunosilent. In APA capsules, high M alginates are preferred as they mask polycations more efficiently through stronger electrostatic interactions.

A surface-to-volume ratio that optimizes implantation efficiency and diffusional transfer processes.

The ability to tailor the capsule towards the host environment. Minimal adhesion or fibrotic overgrowth is tolerable for microcapsules implanted in the peritoneal cavity; however, vascularization is required for subcutaneous transplantation.

A semi-permeable membrane in which the porosity is either too small and prevents diffusion of essential nutrients or oxygen (hypoxia) to the encapsulated cells, or is too large, enabling cells to escape, inducing an immunogenic response, transfection or for cytotoxic species (e.g. inflammatory cytokines such as tumor necrosis factor TNF-α) to enter the microcapsule. Exposure of functionalities that cause a host immune response in the form of attack by humoral and T-cell mediated processes or the formation of a collagenous layer to isolate the ‘foreign’ body. In addition, a rough outer surface of the capsule encourages fibrotic overgrowth, which inhibits diffusion of nutrients and oxygen to the encapsulated cells. Large microcapsules which restrict diffusion of nutrients and oxygen to the encapsulated cells. The surfaceto-volume ratio results in inefficient bioactive molecule release. Physicochemical properties of the outer microcapsule surface that are detrimental to biointegration.

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Biointegration of medical implant materials Desirable: • oxygen and nutrients • signalling biomolecules Undesirable: • stimulus of macrophages, monocytes or fibroblasts which produce antibodies, cytokines or free radicals

Key Encapsulated cell Alginate Calcium ion Cationic polyelectrolyte Anionic polyelectrolyte

Desirable: • cellular waste products • secreted active biomolecules Undesirable: • escaping cells • immunogenic proteins

9.7 Schematic representation of the most common example of a multilayered assembly applied to microencapsulate cells. A summary of the desirable and undesirable processes that can impact on the viability of the cells is shown.

and function of the encapsulated cells. The composition can be altered to vary the mechanical properties and density of the Ca-alginate matrix which can impact on cell proliferation or differentiation, e.g. in the maturation of encapsulated follicles (West et al., 2007). However, the instability of Ca-alginate capsules in vivo due to exchange of crosslinking ions with native ions has led to the increasing use of Ba2+ as the preferred ionic crosslinker. Microcapsules prepared from Ba-alginate hydrogels have been demonstrated to possess improved mechanical properties and an increased capsule lifetime by a factor of two for high G alginates (Mørch et al., 2006), although the ratio of G content needs to be greater than 60%. These microcapsules do not completely prevent the undesirable diffusion of IgG and several cytokines to the cells within the interior of the microcapsule; however, the evidence indicates that the encapsulated islets have sufficient protection from the local environment (Omer et al., 2005). The diffusion of displaced Ca2+ ions within the hydrogel has been reported as having a detrimental impact on the viability of adjacent encapsulated Sertoli cells and the application of

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Glial cell line-derived neurotrophic factor, GDNF Osteocalcin

Schwann

Bone marrow cells (BMC)

Foetal myoblasts

Fischer rat 3T3 fibroblasts

C2C12 myoblasts

HEK293

Crandall–Reece feline kidney (CRFK)

Glial cell line-derived neurotrophic factor, GDNF Therapeutic gene products Total proteins

Glial cell line-derived neurotrophic factor, GDNF Erythropoietin, EPO

Steroids

Follicles

Preosteoblastic (MC3T3-E1)

Secreted biomolecule

Cell system

Purified, intermediate G

Not specified

High G content Low viscosity A range of purity alginates and viscosities Not specified

Purified, high G and both high and low mw RGD-modified alginate† Tyramine-modified alginate for covalent cross-links Non-purified

Purified; G content 55–65%, oxidized and irradiated Filtered High M content

Alginate purity and viscosity

(Ca-Alg)-PLL-Alg

In vitro

In vitro

Striatum

(Ca-Alg)-PLL-Alg (Ca-Alg)-PLL-Alg

Subcutaneously and intraperitoneally

In vitro

Ca-Alg/collagen composite (Ca-Alg)-PLL-Alg

In vitro

(Tyr-Alg) + (Ca-Alg)

In vitro

(Ba-Alg) In vitro

In vitro

(Ca-Alg)

(RGD-Ca-Alg)

In vitro or in vivo

Microcapsule assemblies

Table 9.4 Categories of cells other than pancreatic islets that have been immunoisolated in alginate microcapsules

Abbah et al., 2008

Li et al., 2008

Grandoso et al., 2007

Orive et al., 2005 Ponce et al., 2006

Sakai and Kawakami, 2007 Lee et al., 2009

Evangelista et al., 2007

de Guzman et al., 2008

West et al., 2007

References

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A-aromatase and IGF-1 Secreted alkaline phosphatase (SEAP) reporter enzyme Metabolizing molecules diffusing into capsule Glycosaminoglycan and collagen *

Sertoli cells

Endostatin

Recombinant CHO

Filtered

Purified low viscosity, M rich

Filtered G content not specified G content 69%, 130 K mw Ultrapure, G content 67%, low viscosity

Ultrapure low viscosity G alginate

High purity

Alginate purity and viscosity

Zhang et al., 2008 Zhang et al., 2007

Peritoneal

Marsich et al., 2008 Dusseault et al., 2008

Lin et al., 2008

Luca et al., 2007 Wikström et al., 2008

References

Peritoneal

Peritoneal

In vitro

Ca-Alg+Chitlac‡ (Ca-Alg)-(PLLANB-NOS)-Alg PLL modified with photo-crosslinker to introduce covalent bonds (Ca-Alg)-Chi-Alg (Ca-Alg)-PLL-Alg (Ca removed by chelation) (Ca-Alg)-PLL-Alg (Ca removed by chelation)

Oral administration to stomach

In vitro

Peritoneal

In vitro or in vivo

(Ca-Alg)-Chi-Alg (Ca-Alg)-PLL-Alg

(Ca-Alg)-PLL-Alg Ba-Alg (Ca-Alg)-PLL-Alg (Ba-Alg)-PLL-Alg (Sr-Alg)-PLL-Alg

Microcapsule assemblies

* Study demonstrated retention of immortal cells in microcapsule as proof of concept for similar cell lines that possess risk of malignant transformation. † Modification of alginate to enhance adhesion, proliferation and differentiation. ‡ Chitlac, lactose modified chitosan.

Bone morphologenic protein (BMP)

Embryonic fibroblasts (C3H10T1/2 line)

EL-4 thymoma

Chrondrocytes

E Coli DH5

Human retinal pigment epithelial cell line (ARPE-19)

Secreted biomolecule

Cell system

Table 9.4 Continued

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Ba-alginate microcapsules was reported a successful alternative strategy (Luca et al., 2007). 2 A multilayered self-assembled polyelectrolyte complex (e.g. the APA system) is typically produced by the following sequence: a Ca-alginate hydrogel capsule formed around the cells and then multiple polyelectrolyte layers of opposing charge (e.g. PLL and alginate) are selfassembled through the combination of electrostatic intermolecular forces, hydrogen bonding and polymer flocculation to strengthen the capsule. This assembly remains the basis of the majority of cellular encapsulation matrices. The structure of these capsules is often naively represented as discrete layers; however, evidence has been provided for the formation of polyelectrolyte complexes between alginate and PLL where PLL penetrates the alginate core up to a depth of 30 μm (de Vos et al., 2006). The most significant challenge is to achieve full screening of the exposed charges on the cationic polyelectrolyte, otherwise these will invoke an immune response. PLL inefficiently complexed with an outer alginate layer can potentially remain exposed at the outermost surface of the capsule (Tam et al., 2005). The APA microcapsule offers greater mechanical strength and protection of the encapsulated cells than the Ca-alginate hydrogel capsule; however it must be considered that the longevity of the process, which includes multiple washing steps, can itself be detrimental to the cells. 3 Multilayered assemblies, where the analogues of the APA matrix components are modified to introduce desirable attributes such as substitution of divalent ions or introduction of covalent bonds between biopolymers. The standard APA capsule adopts Ca2+ ions as the crosslinker for the inner capsule. However, the influence of the size of the crosslinking ion (Ca2+, Ba2+, Sr2+) on the structure of the resulting matrix and encapsulated cell viability has been explored for ARPE-19 cells in vitro (Wikström et al., 2008): Sr2+ was found to be unsuitable. The introduction of covalent linkages increased the mechanical strength and reduced the swelling of the capsule matrix in the host environment. The most common modification to the self-assembled LbL matrix is the introduction of covalent bonds between alginate and PLL; for example, those achieved by the introduction of a photocrosslinkable moiety onto the PLL to strengthen the microcapsule (e.g. Dusseault et al., 2008) or the introduction of enzymatically crosslinkable groups (Sakai and Kawakami, 2008). The mechanical properties of the capsule become important in the encapsulation of cells that are required to differentiate and proliferate, and an example of a post-modification strategy is the removal of calcium from the internal hydrogel, to encourage cell growth and active biomolecule

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production (Zhang et al., 2008). The modification of alginate through the introduction of peptide moieties (Section 9.2.4) is a common approach to improving the adhesion of encapsulated cells that are anchorage dependent for survival, e.g. osteoblasts (Evangelista et al., 2007). An extension of this approach was demonstrated in the covalent modification of alginate by YIGSR peptides prior to encapsulation of neurites, followed by the subsequent modification of the capsule via adsorption of laminin (Dhoot et al., 2004). Adhesion of the neurites was promoted. Post-modification of the encapsulated cell assembly has been attempted to improve integration with its target environment and strategies include PEGylation (Zhang et al., 2008). Surface functional molecules may also be coupled to the assembly post-encapsulation to improve immunosilence (Leung et al., 2008). An emerging initiative in the microencapsulation of cells is to encapsulate active biomolecules simultaneously with cells, with the objective of promoting a subsequent cellular response in vivo. Bioactive molecules, such as basement membrane extract BD MatrigelTM (de Guzman et al., 2008) to promote proliferation, have been incorporated with the alginate prior to formation of the Ca-alginate microcapsule. In a separate approach, extracellular matrix proteins, such as collagen and laminin, have been added during the assembly of the polycationic membrane (Cui et al., 2006). Finally, there are a number of recent studies that demonstrate that researchers are thinking ‘beyond the sphere’ in cell encapsulation by tailoring the scaffold geometry and matrix properties to suit their specific applications. There are applications where the cells themselves are ‘delivered’ as they migrate from a scaffold that is designed to co-deliver inductive molecules (Hill et al., 2006). An innovative approach is represented in the encapsulation of cells within fibres which may also contain growth factors (Wan et al., 2004). These fibres can then be self-assembled into defined and patterned structures.

9.4

Future trends

At this point the influence of the structure of alginate (M/G ratio, chain length and role of naturally derived impurities) has been well characterized. The process of assembling three-dimensional matrices through crosslinking of alginate molecules is also well understood and can be applied to control mechanical properties and stability. Modification of the assemblies is emerging as the popular route for optimizing biointegration of these matrices in their host environment. Research encompasses two approaches: the modification of alginate prior to capsule assembly and post-modification of assembled matrices. As the process of assembling alginate-based matrices for drug delivery is refined, attention will turn to tailoring the release profiles of the bioactive

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species. The nature of the interaction between the drug molecules and the scaffold matrix will define these profiles. To date, knowledge of the chemical properties of the drug molecule has been poorly explored as the delivery vehicle is designed. Introduction of functionalities that can be manipulated through change of environment or time-dependent processes offer new strategies, and already reports of pH responsive systems through the modification of alginate have appeared (Chan et al., 2008). Finally, the majority of assemblies addressed here have resulted in spherical capsules. Drug delivery is not restricted to matrices of this geometry and while the surface area is reduced in other three-dimensional geometres, there are advantages offered by drug delivery scaffolds such as self-assembled fibres (Wan et al., 2004) and semi-interpenetrating networks (Matricardi et al., 2008).

9.5

Acknowledgement

The authors are grateful to the Australian Research Council (Grant No DP0557475) for their support of this work.

9.6

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vold, i m n, kristiansen, k a, christensen, b e (2006) ‘A study of the chain stiffness and extension of alginates, in vitro epimerized alginates, and periodate-oxidized alginates using size exclusion chromatography combined with light scattering and viscosity detectors’ Biomacromolecules, 7(7), 2136–2146. wan, a c a, yim, e k f, liao, i-c, le visage, c, leong, k (2004) ‘Encapsulation of biologics in self-assembled fibres as biostructured units for tissue engineering’ Journal of Biomedical Materials Research, 71A, 586–595. webber, r e, schull, k r (2004) ‘Strain dependence of the viscoelastic properties of alginate hydrogels’ Macromolecules, 37(16), 6153–6160. west, e r, xu, m, woodruff, t k, shea, l d (2007) ‘Physical properties of alginate hydrogels and their effects on in vitro follicle development’ Biomaterials, 28, 4439–4448. whistler, r l, wolfrom, m l (1962) Methods in Carbohydrate Chemistry, Vol I: Analysis and Preparation of Sugars. 2nd edition, New York and London, Academic Press Inc. wikström, j, elomaa, m, syväjärvi, h, kuokkannen, j, yliperttula, m, honkakoski, p, urtti, a (2008) ‘Alginate-based microencapsulation of retinal pigment epithelial cell line for cell therapy’ Biomaterials, 29, 869–876. wilson, j t, chaikof, e l (2008) ‘Challenges and emerging technologies in the immunoisolation of cells and tissues’ Advanced Drug Delivery Reviews, 60, 124–145. xu, j b, bartley, j p, johnson, r a (2003a) ‘Preparation and characterization of alginate hydrogel membranes crosslinked using a water-soluble carbodiimide’ Journal of Applied Polymer Science, 90, 747–753. xu, j b, bartley, j p, johnson, r a (2003b) ‘Preparation and characterization of alginate–carrageenan hydrogel films crosslinked using a water-soluble carbodiimide (WSC)’ Journal of Membrane Science, 218, 131–146. yeom, c k, lee, k h (1998) ‘Characterization of sodium alginate membrane crosslinked with gluteraldehyde in pervaporation separation’ Journal of Applied Polymer Science, 67, 209–219. zhang, y, wang, w, xie, y, yu, w, lv, g, guo, x, xiong, y, ma, x (2007) ‘Optimization of microencapsulated recombinant CHO cell growth, endostatin production and stability of microcapsule in vivo’ Journal of Biomedical Materials Research B: Applied Biomaterials, 84B, 79–88. zhang, w-j, li, b-g, zhang, c, xie, x-h, tang, t t (2008) ‘Biocompatability and membrane strength of C3H10T1/2 cell-loaded alginate-based microcapsules’ Cytotherapy, 10, 90–97. zhu, h, ji, j, lin, r, gao, c, feng, l, shen, j (2002) ‘Surface engineering of poly(dl-lactic acid) by entrapment of alginate–amino acid derivatives for promotion of chondrogenesis’ Biomaterials, 23, 3141–3148. zimmermann, u, klock, g, federlin, k, hannig, k, kowalski, m, bretzel, r g, horcher, a, entenmann, h, sieber, u, zekorn, t (1992) ‘Production of mitogen-contamination free alginates with variable ratios of mannuronic acid to guluronic acid by free flow electrophoresis’ Electrophoresis, 13(5), 269–274. zimmermann, h, zimmermann, d, reuss, r, feilen, p j, manz, b, katsen, a, weber, m, ihmig, f r, ehrhart, f, gebner, p, behringer, m, steinbach, a, wegner, l h, sukhorukov, v l, vasquez, j a, schneider, s, weber, m m, vole, f, wolf, r, zimmermann, u (2005) ‘Towards a medically approved technology for alginate-based microcapsules allowing for long-term immunoisolated transplantation’ Journal of Materials Science Materials in Medicine, 16, 491–501.

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10 Functionalised nanoparticles for targeted drug delivery S. M A N J U and K. S R E E N I VA S A N, Sree Chitra Tirunal Institute for Medical Sciences and Technology, India

Abstract: Nanocarriers composed of liposomes, micelles, polymeric nanoparticles and others have shown tremendous opportunities in the field of targeted drug delivery, especially in cancer therapy. Functionalisation of nanomaterials through simultaneous assembling of chemical moieties has been a strategy of wide interest, to acquire properties such as longevity in circulation, site specificity and stimuli sensitivity. Imparting multifunctionality to nanocarriers controls their biological interaction in a desired fashion and enhances the efficacy of therapy and diagnostic protocols. Here, we attempt to review the application of various nanocarrier systems for targeted drug delivery and current strategies for the development of multifunctionality on nanocarrier systems. Key words: nanocarriers, drug delivery, micelle, liposomes, polymeric nanoparticle, gold nanoparticle, magnetic nanoparticle, hyperthermia, photodynamic therapy.

10.1

Introduction

Nanobiotechnology is a multidisciplinary field that covers a vast and diverse array of technologies coming from engineering, physics, chemistry and biology. It is the combination of these fields that has led to the birth of a new generation of materials and methods of making them. Nanomaterials, which measure 1 nm to 100 nm, allow unique interaction with biological systems at the molecular level. Among the various approaches for exploiting developments in nanotechnology for biomedical application, nanoparticulate carriers offer some unique advantages as delivery, sensing and image enhancement agents.1,2 They can engineer important advances in the detection, diagnosis and therapy of human cancers. We know that many effective bioactive agents used for pharmacotherapy exhibit side effects that limit their clinical application, so it is important to achieve selectivity in the delivery of drug molecules to target areas, in order to enhance therapeutic potential and minimise side effects. For example, cytotoxic compounds used in cancer therapy can kill cancer cells as well as normal cells. The use of pharmaceutical nanocarriers is designed to overcome most of 267 © Woodhead Publishing Limited, 2010

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these limitations.3 Even though research efforts in this area have resulted in enhanced in vitro efficacy of many drugs, both in pharmaceutical research and a clinical setting, researchers are still engaged in designing improved nanocarriers for effective and targeted delivery applications. In general, requirements in the design of nanodrug delivery systems include: (a)

design carrier systems that can incorporate different types of therapeutic agents in sufficient doses, (b) recognise disease-specific ligands that can be conjugated to drug carrier systems and achieve targeted drug therapy, (c) protect drug molecules from degradation in the body prior to their delivery at the required sites, (d) develop drug-carrier system that can release the drug at the target site at a desired or controllable rate for the duration necessary to elicit the desired pharmacological response, (e) develop drug carrier systems that are biocompatible and biodegradable so that these can be used safely in human, and (f) achieve effective intracellular drug delivery for those therapeutic agents whose receptor or site of action is intracellular. Nanocarrier systems are formulated from a variety of materials having various chemical compositions, and are engineered to carry a number of bioactive molecules in a controlled and targeted manner, making them efficient drug delivery vehicles.4 Commonly used nanocarriers in drug delivery application include liposomes,5 micelles, nanocapsules, polymeric nanoparticles, solid lipid nanoparticles, metallic nanoparticles (gold/magnetic nanoparticle/quantum dots) and others (Fig. 10.1).5–8 Some of the examples of commonly defined nanocarrier systems and their medical applications are shown in the Table 10.1. Generally, the selection of material for developing nancarriers is mainly dictated by the desired diagnostic or therapeutic effect, administration route and type of pay load. It offers considerable opportunities to improve cancer therapy by exploring the unique inherent properties of a solid tumor combined with new approaches such as local heating or reactive oxygen generation. Nanocarriers, depending on the reticular requirement, demand surface modification.10–13 Surface modified nanocarriers exhibit advantageous properties, such as prolonged circulation in the blood, demanded biodistribution, passive or active targeting to the pathological tissue site, responsiveness to local physiological stimuli (pH or temperature changes in the tissue site with pathological condition), and the ability to serve as imaging or contrast agents for various imaging systems. This review focuses mainly on different targeting methods and recent progress in the functionalization of nanocarriers for improving cancer therapy, including both imaging and site-specific targeting of drug molecules.

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10–3 m (mm) Spleen cut-off 200 nm

Cells Lymphocytes Erythrocytes DNA

10–6 m (μm) Lipid bilayer

Proteins

Dendrimers Micelle Emulsion Nanoparticulate Liposomes 3–5 nm 5–10 nm water in oil systems 100–150 nm Drug 100–150 nm 20–150 nm 10 m (nm) encapsulated or attached

Quanfum dots

–9

Bucky balls

Nanoparticles used in drug delivery Small molecules and atoms

10–12 m (pm)

10.1 Nanoparticle systems in drug delivery (adapted with permission from © 2007 Elsevier9).

10.2

Drug targeting

Drug targeting is a potential approach by which distribution of a drug in an organism is contrived in a manner such that its major fraction interacts exclusively with the target tissue at the cellular or sub-cellular level. Theoretically, drug targeting can improve the outcome of chemotherapy by means of one or both of the following processes: (a)

(b)

By allowing the maximum fraction of the delivered drug molecule to react exclusively with diseased cells without adverse effect to the normal cells. By allowing preferential distribution of the drug to the diseased/cancerous cells.

10.2.1 Barriers to drug targeting The main hurdles in the field of drug targeting include physiological barriers and biochemical challenges to achieving target specificity. They also include the selection of appropriate techniques for the conjugation of the targeting ligand to the nanocarriers. The challenges in drug targeting include targeting drugs to specific sites and imparting longevity in the blood to enhance bioavailability at the site of action. For intravenous administration of nanocarriers, the main barrier is that of the vascular endothelium, basement membrane14 and plasma proteins. Plasma proteins have the ability to

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Table 10.1 Commonly used nanoparticles and their medical applications Nanoparticle

Examples

Medical application

Metal nanoparticles

Quantum dots Gold nanoparticles Gold nanorods Gold nanoshells Gold nanocages Magnetic nanoparticles

Diagnostics Biosensor Molecular imaging Drug delivery

Nanotubes, nanowires

Carbon nanotubes

Biomolecular sensing Delivery of vaccines or proteins

Dendrimers

Poly(amido amine)

Drug carriers Imaging agent Gene delivery

Liposomes

PEGylated immunoliposomes

Drug delivery Gene encoding

Polymer micelles

Doxorubicin conjugated to poly(ethylene glycol)-poly (α,β-aspartic acid) PEG-PAsp(DOX)

Drug delivery of water-insoluble drugs

Ceramic nanoparticles

Silica-based nanoparticles entrapping photosensitising anticancer drug, 2-devinyl2-(1-hexyloxyethyl) pyropheophorbide

Drug delivery

Polymeric nanoparticles

PLGA (poly(D,L-lactic-coglycolic acid) Poly(lactic) acid (PLA)– polyglycolic acid (PGA)

Drug delivery Protein delivery Gene expression vector

Polysaccharide nanoparticles

Cellular nanocrystals

Targeted delivery Bioimaging

Magnetic nanoparticles

Superparamagnetic iron oxide

Magnetic resonance imaging, contrast agent

Bionanoparticleprotein-based nanosystems

Ferritin, viruses and virus-like particles Heat shock protein cages

Gene delivery Bioimaging Drug delivery Vaccine development

adversely affect the biodistribution of drug carriers introduced in the blood stream. The in vivo biodistribution and opsonisation of nanosystems in blood circulation are governed by their size and surface characteristics, such as zeta potential and hydrophilicity/hydrophobicity, so one has to be very critical in developing targeted nanocarriers and should avoid opsonisation

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and subsequent recognition by reticuloendothelial cells (RES). Another barrier is that of the extra cellular matrix, which has to be crossed to access the target cells in a tissue. On dealing with intracellular delivery, there are additional barriers to be crossed to allow internalisation of the systems into the specific cells.15 Also, there are a number of endocytic pathways for the cellular entry of nanosystems. Researchers have to be aware of the particular mechanisms for a better demonstration of the functionalisation of nanocarriers.The nuclear membrane creates another formidable barrier for drugs such as oligonucleotides, plasmid DNA and other low molecular weight drugs, whose site of action is located in the nucleus of a cell. During recent years, a number of cellular and molecular targets have emerged in the field of drug delivery, but poor bioavailability of the drug in the target tissue is a real problem in clinical practice due to these barriers. Thus, to develop competent targeted systems, it is necessary to successfully overcome most of the physiological barriers and deliver the drug in the target site at an optimum therapeutic level, for the required time period to elicit pharmacological action. Various techniques have been devised for the conjugation of targeting ligands onto nanocarrier systems, which include covalent and non-covalent conjugation. These approaches direct them to their target site and give them the right orientation for binding the target molecule.16 The conjugation mechanism should not adversely affect the integrity of the nanocarriers and should give a better interface to enhance the biological activity. To this end, various strategies have been developed for targeting drugs to the required tissue site in the body by the proper design of nanocarriers. The various approaches for targeting drugs in chemotherapy are shown in the Fig. 10.2. Drug targeting is mainly classified into passive targeting and active targeting. Site-specific drug targeting

Passive targeting

EPR effect

Active targeting

Tumour environment

Carbohydrate interaction

Ligand–receptor interaction

Antibody– antigen interaction

Direct local delivery

10.2 Schematic representation: various approaches for drug targeting (EPR, enhanced permeability and retention).

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6–7 nm ECs

200–800 nm

Lumen

ECM Normal tissue, intact vasculature

Tumour tissue site, leaky vasculature

10.3 Pathological difference between normal and tumour tissue leads to better accumulation of drug in the tumour site.

10.2.2 Passive targeting Solid tumours present favourable conditions for the preferential accumulation of a variety of nanocarriers. The rapid growth of a solid tumour results in altered physiology at the tumour site, which leads to leaky and defective vasculature (gaps ~ 600 nm, Fig. 10.3) and poor lymphatic drainage. The increased vascular permeability coupled with impaired lymphatic drainage in the tumour, leads to an enhanced permeability and retention effect (EPR effect) which allows extravasations of nanocarriers and selective localisation in the inflamed tissue.17,18 Oral administration of polymeric nanocarriers could, as an example, be selectively targeted to the inflamed colonic mucosa in inflammatory bowel disease.19 Similarly, in some critical inflammatory conditions, the blood–brain barrier can be crossed to access the target sites for brain delivery.20 For enhancing the passive targeting mechanism, the physicochemical factors of the nanocarriers such as size, surface charge and surface hydrophobicity also play critical roles. Particles <100 nm can pass through the fenestrations in the liver endothelium and the sieve plates of sinusoids to localise in the spleen and bone marrow. This natural tendency of nanosystems to localise in the reticuloendothelial system (RES) presents an excellent opportunity for the passive targeting of drugs to the macrophages present in the liver and the spleen. Although the RES uptake of the colloidal-sized nanocarriers holds promise for passive targeting of many drugs, it becomes a major hurdle for drugs whose site of action is located in tissues other than the RES. A variety of methods have been explored to manipulate rapid opsonisation and subsequent RES uptake of nanocarriers, and to make them long circulating for effective targeting of drugs to specific tissues site. Hydrophilic polymers play an important role in these aspects and enhance longevity in the blood via imparting steric stabilisation.

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One of the widely used hydrophilic polymers is poly(ethylene glycol), (PEG), which sterically hinders the interactions of blood components with tissue surfaces and reduces the binding of plasma proteins.21–24 This prevents drug carrier interaction with opsonins and delays their fast capture by the RES.25 Here, PEG opposes the opsonisation mechanisms by shielding the surface charge on the nanocarriers and by imparting increased surface hydrophilicity. Increased surface hydrophilicity along with the steric hindrance effect of PEG leads to enhanced repulsive interaction between the functionalised nanocarrier system and blood components.26 Some of the hydrophilic and hydrophobic polymers can also be used as alternative steric protectors for these nanocarrier systems;27 for example, single end lipid modified poly(acrylamide) and poly(vinylpyrrolidone), together with amphiphilic polymers such as phosphatidyl polyglycerols, poly(acryloyl morpholine) (PAcM), phospholipid-modified poly(2-methyl-2-oxazoline) or poly(2-ethyl-2-oxazoline). Fluorouracil containing dendrimer nanoparticles modified with PEG have demonstrated better drug retention and less haemolytic activity.28 Polycyanoacrylate nanoparticles modified using PEG showed increased longevity in the circulation, even allowing for their diffusion into the brain tissue.29 Nanocarriers based on a poly(lactic–glycolic acid) PEG (PLAGA-PEG) copolymer act as a long-circulating system having an insoluble (solid) PLAGA core on which a water-soluble PEG shell is covalently attached. The PLAGA microspheres modified by adsorption of the polylysine-PEG copolymer help to attain a dramatic decrease of plasma protein adsorption.30 Also, polyvinyl alcohol has been successfully used as a liposome steric protectors.31–36 In addition to the decrease of particle uptake by resident macrophages in liver, coating with some specific copolymers can redirect the injected nanoparticles to other organs (thus, a coating of 60 nm polystyrene latex with Poloxamer 407 results in increased particle accumulation in bone marrow)37 and the PEG imparts long circulating nature to the polystyrene latex. Long circulating liposomes containing various anticancer agents, such as doxorubicine, arabinofuranosylcytosine, adriamycin, and vincristin were prepared. PEG-liposome incorporated doxorubicine has already demonstrated very good clinical results.38–40 Recently, several studies have emerged in the field of superparamagnetic nanoparticles in drug delivery and diagnostic imaging.41 Similarly to other nanocarriers, PEG-modified magnetite nanospheres have shown an increased colloidal stability and improved localisation in lymph nodes.42 Also, PEG can be grafted onto the surface of gold particles using mercaptosilanes, and enhances the reduction in protein adsorption and platelet adhesion.43 Hence, the nanocarriers modified with protecting polymers are more effective in clinical practice by increasing circulation time with decreased RES uptake, leading to effective accumulation in the tumour via the EPR effect.

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10.2.3 Active targeting The extent of passive targeting approaches is limited and therefore, tremendous efforts have been directed towards the development of active approaches by conjugating target-specific ligand molecules on the nanocarriers that offer preferential accumulation in the tumour bearing organs. The receptors and surface-bound antigens are expressed uniquely/differentially in a higher manner in diseased cells as compared to normal cells, and explore the possibilities of active targeting. The presence of specific receptors on cell membranes helps in active targeting by ligand-specific interaction of nanocarriers with cells.44 Various biological molecules, such as receptors, antibodies and proteins, have been used extensively to target specific cells. These ligand molecules are bound to the surface of nanocarriers via chemical coupling and take advantages of ester or amide linkags. Table 10.2 shows some important ligand molecules commonly used in active targeting. A proper understanding of these systems explores the possibility Table 10.2 Protein/ligand molecules commonly used in active targeting and its application Protein or ligand

Functional activity

Transferrin

Widely applied as targeting ligand in the active targeting of anticancer agents, proteins and genes to primary proliferating cells via transferring receptor Structurally similar to transferrin, globular multifunctional protein with antimicrobial activity. Promotes proliferation and differentiation of cells and important for normal wound healing A hormone that regulates blood glucose levels Promotes neuritis outgrowth and neural cell survival Principal carrier for copper in plasma, which plays an important role in iron haemostasis and is also an effective anti-oxidant for a variety of free radicals High water solubility, no toxicity, usefulness as a plasma expander, non-immunogenic, non-antigenic properties A cross-linked protein in the extracellular matrix that provides elasticity of many tissues The major serum protein, binds a wide variety of lipophilic compounds including steroids, etc. Membrane-permeating peptide, enhances cellular delivery Increases cell spreading, differentiation, enhances intracellular delivery Preferentially targets cancer cells, poorly immunogenic, folate receptor facilitates internalisation of particle

Lactoferrin Transforming growth factor (TGF-α) Insulin Nerve growth factor Ceruloplasmin

Pullulan

Elastin Albumin TAT peptide RGD peptide Folic acid

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of drug targeting in a more specific manner. For example, liposomes bind sufficient numbers of antibody molecules to their surfaces without compromising the liposome integrity or antibody affinity and specificity. One of the routine methods for antibody coupling to liposomes includes protein covalent binding to the reactive group on the liposome membrane. It is well reviewed in the literature that chemical modification of a non-membrane hydrophilic protein (such as an antibody) with a hydrophobic reagent (such as phospholipids residue) increases the affinity of the modified protein towards the liposomal membrane. It has been shown that hydrophobic groups can be easily incorporated into proteins by treating them with long chain fatty acid chlorides45 or activated phospholipids (e.g. oxidised phosphatidyl inositol,46 or N-glutaryl phosphatidyl ethanolamine).47 With this flexibility, numerous methods have been developed during the past decades for coupling antibodies to liposomes. Nanoparticles made of gelatin and human serum albumin have also been modified, with HER2 receptor specific antibody (Trastuzumab) via an avidin–biotin system. The study showed that these surface-modified nanoparticles were effectively endocytosed by HER2 over-expressing cells. In another study, antibody-specific for CD14 and prostate-specific membrane antigen were used to modify the surface of dendrimer nanoparticles and showed specific binding to corresponding antigen over-expressing cells.48 Further developments involved the combination of longevity and target specificity in one preparation by simultaneous co-incorporation of antibody and PEG onto the liposome. A schematic representation of multifunctional surfaces in different nanocarrier systems is shown in Fig. 10.4. However, the careful selection of a protective polymer/targeting moiety ratio also provides potential efficacy in drug targeting. In this case, PEG compensates for the effect of an increased rate of elimination from the blood circulation of the ligand-conjugated species. Longevity of the specific ligand-bearing nanocarriers may allow for preferential accumulation, even in targets with diminished blood flow or with low surface antigen concentrations.50 To achieve better selective targeting by PEG-coated liposomes or other particulates, targeting ligands can be attached to nanocarriers via the PEG spacer arm, so that the ligand is extended outside of the dense PEG brush, excluding steric hindrances for its binding to the target receptors. Recently, Favrokhzad et al. have demonstrated the use of aptamer and nucleic acid, which specifically recognise prostate membrane antigen on prostate cancer cells and provide additional opportunities for active targeting.51,52 Intracellular delivery and subcellular distribution Once the nanocarriers are delivered to the target cells, they may be internalised by specific or nonspecific cell penetrating strategies, endocytosis. In

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Liposomal systems

Solid biodegradable polymer nanoparticle

Phospholipid bilayer Multifunctional surface

Biodegradable core

PEG

10–1000 nm

100–1000 nm

(c)

Interior contrast agent Dendrimer (e.g. iron oxide) Exterior contrast agent (e.g. Gd-DTPA-lipid) Drug

Dendrimer core

Antibody Other target ligand PEG

10–100 nm

10.4 Schematic representation of multifunctional surfaces in various nanocarrier systems (Gd-DTPA, gadolinium diethylenetriaminepentaacetic acid) (adapted with permission from © 2007 American Association of Pharmaceutical Scientists49).

nonspecific uptake of nanocarriers, cell membrane envelops the nanocarriers to form endosome. However, specific cellular uptake can occur through receptor-mediated endocytosis. These receptors are recycled back to the cell surface on dissociation of the nanocarrier receptor complex. For example, in the case of targeting of folate-conjugated nanocarrier, folate receptors are over-expressed in many tumour cells. Figure 10.5 shows the folatemediated drug targeting of nanoparticles and intracellular uptake by endocytosis. Liposomal daunorubicin54 and doxorubicin55 have been delivered into various tumour cells via folate receptor and demonstrated increased cytotoxicity. Also, folate-targeted liposomes have been used as delivery vehicles for boron neutron capture therapy.56,57 Folate was also attached to the surface of cyanoacrylate-based nanoparticles via activated PEG blocks.58 Similarly, PEG–polycaprolactone-based particles were surface modified with folate and, after being loaded with paclitaxel, demonstrated increased cytotoxicity.59 Transferring (Tf) receptor (TfR) is over-expressed on the surface of many other tumour cells. And hence antibodies against TfR, as well as Tf itself, act as effective targeting for targeting various nanoparticular drug carriers, including liposomes to tumours and inside tumour cells.60

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Folate receptors

Receptor-mediated endocytosis

Nucleus

Folate-mediated nanoparticles

Endocytosis

Intracellular pH-dependent drug release

Drug-loaded core

10.5 Folate-mediated drug targeting of nanoparticle and intracellular uptake by endocytosis (adapted with permission from © 2008 Elsevier53).

Recent studies have involved the coupling of Tf to PEG on PEGylated liposomes in order to combine longevity and targetability for drug delivery into solid tumours.61 Tf-mediated liposome delivery has also been successfully used for brain targeting. Stimuli-responsive nanocarriers Stimuli-responsive nanocarriers can help to address some of the systemic and intracellular delivery barriers and offer interesting opportunities for drug and gene delivery. Triggers in each diseased site are unique (e.g. increased temperature, decreased pH and hypoxia) and hence a proper understanding of the difference between normal and pathological tissues helps to explore site-specific targeting using stimuli-responsive nanocarriers. pH-responsive nanocarrier systems Nanocarriers constructed from stimuli-responsive polymers have been used effectively in passive/active cancer therapy. The pH profile of pathological tissues having inflammation, infection, and cancer, is significantly

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different from that of normal tissue.62 For example, most solid tumours have lower extracellular pH (<6.5) than the surrounding normal tissues (pH 7.4). Also pH is compartmentalised in tumour tissue into an intracellular component (which is similar in tumour and normal tissue) and an extracellular component, which is relatively acidic in tumours.63 This gives rise to a cellular trans-membrane pH gradient difference between normal tissue and tumour tissue, which may also be exploited for the delivery of drugs to the tumour site.64 Polymeric compounds with pH-sensitive bonds can be used to produce stimuli-responsive drug delivery systems. Physical properties, such as swelling/deswelling, particle disruption, and aggregation, of stimuli-responsive nanocarriers change in response to changes in environmental conditions, alter the interactions of the nanocarriers with the cells, and trigger the drug release from slow to fast at the tumour site. For example, the pH-sensitive poly(β-amino ester)s (PbAE) constitute a novel class of biodegradable cationic polymers for development of site-specific drug and gene delivery systems. In the acidic microenvironment of the tumour, PbAE undergoes rapid dissolution and quickly releases its contents.65 Water-soluble polymeric drug carriers based on a copolymer of N-(2-hydroxypropyl methacrylamide) (HPMA) or a linear polymer consisting of PEG blocks have been developed as drug delivery systems capable of delivering drugs to model tumours or tumour cells. Research based on pullulan acetate conjugated with sulfadimethoxin (pKa 6.1) has been developed to prepared pH-sensitive and self-assembled hydrogel nanoparticles that demonstrate enhanced adriamycin release in response to lower pH and increased cytotoxicity. A summary of different types of pH-responsive nanoparticulate systems is shown in the Table 10.3. Temperature-responsive nanocarrier systems The use of hyperthermia as an adjunct to radiation or chemotherapy of various types of solid tumours has become an area of active investigation for the past 20 years. It is well established that tumour cells are more sensitive to heat-induced damage than normal cells. Recently, a majority of clinical studies of hyperthermia have used superparamagnetic iron oxidecontaining liposomes or nanoparticles.66,67 The liposomes and nanoparticles provide a method for intracellular delivery and localisation of the iron oxide particles. Unlike external probes that can heat the surrounding normal tissues, the magnetic nanoparticle hyperthermia is appealing because it offers a way to ensure that only the intended target is heated. Thermally responsive polymeric carriers, in combination with hyperthermia, achieve a significant increase in delivery to solid tumours compared with the same polymers without hyperthermia. Meyer et al. have demonstrated that

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pH-sensitive micelles

Doxorubicin

Doxorubicin

Self-assembling amphiphilic block copolymers, poly(ethylene glycol)poly(aspartate hydrazone-adriamycin)

Folate–poly(ethylene glycol) poly(aspartate hydrazone-adriamycin)

Doxorubicin

Poly(N-isopropylacrylamide)

Insulin

Higher drug release under endosomal/ lysosomal low pH conditions (5–5.5) Hydrazone bond cleaved intracellularly at low pH (5–6) and releases the drug

Insulin released in a pH-dependent fashion Cationic polymer induces membrane fusion at endosomal pH Induces a coil-toglobule transition at acidic pH and is exploited to destabilise the intracellular vesicle membrane

Rapid sol-to-gel transition with change in pH

Paclitaxel

Plasmid DNA

Poly(epsilon-caprolactone-co-lactide)poly(ethylene glycol)-poly(epsiloncaprolactone-co-lactide) with sulfamethazine oligomer Poly(N-isopropylacrylamide-cobutylmethacrylate-co-acrylic acid)

pH-sensitive polymeric

Property

Drug or gene

N-Ac-poly(l-histidine)-graft-poly(l-lysine)

Stimuli-responsive polymer or lipid

Nanocarrier

Showed significantly higher antitumour activity in C-26bearing mice These multifunctional micelles enhanced the KB cellular uptake and increased cell kill

Showed improved anticancer activity in murine tumour models

Good antitumour effect in melanoma tumour-bearing mice System has potential as a polypeptide drug carrier Showed higher transfection efficacy in 293T cells

Therapeutic outcome

Table 10.3 Illustrative examples of stimuli-sensitive nanocarriers for drug and gene delivery (adapted with permission from © 2008 Elsevier53)

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pH-sensitive dendrimers

pH-sensitive liposomes

Nanocarrier

Chlorambucil Doxorubicin

Drug linked to polyamidoamine dendrimer via pH-sensitive linker

Antisense oligonucleotides

Mastoparan

Plasmid DNA

Doxorubicin

Polyamidoamine dendrimer

Anionic liposomes containing phosphatidylethanolamine (PE)

Dioleoylphosphatidylethanolamine liposomes with PEG via disulfide linkage Histidine-modified galactosylated cholesterol derivative – cationic liposome Transferrin-modified liposomes (Tf-L) with a pH-sensitive fusogenic peptide (GALA)

Drug released in a pH-dependent fashion Photochemical mediated internalisation of the carrier

Low pH allows fusion of endosomal membranes and destabilisation of the endosomes

Selective delivery to mitochondria

Enhanced intracellular drug delivery Cytoplasmic delivery

Drug release faster at low pH and able to target intracellularly Micelles destabilised in the pH range of 7.2–6.6

Doxorubicin

Folate–poly(N-isopropylacrylamide-coN,N-dimethylacrylamide-co-2aminoethyl methacrylate)-b-poly(10undecenoic acid) Mixed micelles – PLLA/PEG block copolymer with polyHis/PEG Doxorubicin

Property

Drug or gene

Stimuli-responsive polymer or lipid

Table 10.3 Continued

Improved cytotoxicity on Ca9-22 cells

Effective cytotoxicity on B-lymphoma cells Increased gene transfection to hepatocytes Selective mitochondrial targeting for cancer therapy Effective in the treatment of viral infections, cancer or inflammatory diseases

Showed effective MCF-7 cell inhibition and uptake in vitro

Significantly enhanced KB cell growth inhibition

Therapeutic outcome

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Temperaturesensitive liposomes Calcein

Dioleoylphosphatidylethanolamine vesicles bearing poly(N-isopropylacrylamide) Poloxamer F127 containing liposomes

pGFP Plasmid DNA

pH-sensitive TAT-modified PEGylated liposomes

Thiopolycation PESC

Lucifer yellow iodoacetamide

Doxorubicin

Doxorubicin

DPPC:MPPC:DSPE-PEG-2000

Poly(N-isopropylacrylamide-bbutylmethacrylate

Intracellular delivery of pGFP and liposomes Thiopolyplexes releases DNA in reductive environment

Low temperaturesensitive liposomes that trigger complete release in 39–40 °C Complete drug release achieved at 40 °C 90% release achieved at 42 °C

Enhanced drug release in response to temperature fluctuation

Rapid sol-to-gel transition with change in temperature Temperature sensitivity was observed at 42 °C

Paclitaxel

Poly(epsilon-caprolactone-co-lactide)poly(ethylene glycol)-poly(epsiloncaprolactone-co-lactide) with sulfamethazine oligomer Multi-block copolymers

Temperaturesensitive polymeric nanoparticles Doxorubicin

Property

Drug or gene

Stimuli-responsive polymer or lipid

Nanocarrier

Efficient gene transfection

Showed 2.5 fold increase in fluorescence in CT-26 tumourbearing mice Effective tumour therapy possible

Increased therapeutic efficacy in FaDu human tumour xenografts

Showed good antitumour effect in melanoma-bearing mice Showed enhanced activity against Lewis lung carcinoma cells Improved cytotoxicity

Therapeutic outcome

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polymeric nanoparticles that undergo a lower critical solution temperature (LCST) could be designed to be systemically soluble when injected in vivo, but to become insoluble and accumulate in locally heated regions. Examples of polymer systems that exhibit LCST include poly(N-isopropylacrylamide) (poly(NIPAAm)) and elastin-like polypeptides (ELPs). NIPAAm is known to exhibit phase-transition at 32 °C, termed LCST.68 It shows a very sharp change in hydrophilicity/hydrophobicity at a very narrow temperature region around 32 °C. In addition, the LCST of NIPAAm can be easily modified to get above 37 °C using copolymers with varying hydrophilicity or hydrophobicity. PEG is the commonly used hydrophilic segment of the copolymers forming the micelles, as well as for the coating of other colloidal nanocarriers, because of its biocompatibility and good ‘stealth’ properties.69 The micelle’s hydrophobic inner core offers an important control point for temperature-sensitive drug release. Both inner cores and outer shell polymer chemistries have been investigated to modify the temperature-responsive behaviour of micelles for specific drug delivery.70 Cholesteryl end-capped temperature-sensitive amphiphilic polymers were synthesised from hydroxyl-terminated random poly(N-isopropylacrylamide-co-N, N-dimethylacrylamide) (NIPAAm-co-DMAAm) and amineterminated NIPAAm conjugated with cholic acid. These polymers were shown to have LCSTs of 37.7–38.2 °C and 31.5 °C, respectively.71 In another study, amphiphilic NIPAAm-grafted-polyphosphazene (NIPAAm-g-PPP) was synthesised by stepwise substitution of chlorine atoms on polymer backbones with aminoterminated NIPAAm oligomers and ethyl glycinate (GlyEt); the LCST was found to be 30 °C in water. An amphiphilic thermosensitive nanocarrier was prepared from N-(2-hydroxypropyl) methacrylamide lactate and PEG. This was shown to be a promising delivery system for the parenteral administration of paclitaxel.72 Researchers have investigated the modification of liposomes with NIPAAm copolymers to obtain liposomes with temperature sensitivity.73 Han et al.74 investigated the surface modification of liposomes by using poly(N-isopropylacrylamide-co-acrylamide) (NIPAAm-AAM) and PEG. The release of DOX from the NIPAAmAAM/PEG modified liposomes increased around the transition temperature of the polymer. In addition, modified liposomes were found to be stable in the serum, compared with unmodified liposomes, suggesting that NIPAAm-AAM/PEG modified liposomes are suitable for targeted-drug delivery. Some of the temperature-sensitive nanocarrier systems are shown in the Table 10.3.

10.3

Multifunctional nanocarrier systems

Even though various organic materials have been explored for drug delivery using passive targeting or active targeting with a recognition moiety or

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physical stimulus, they have some limitations in terms of chemical or mechanical stability, swelling, susceptibility to microbiological attack, and so on.75 Because of these disadvantages of organic nanoparticles, researchers have also concentrated on inorganic systems having simultaneous imaging and drug delivery capabilities.76 It means that beyond the ordinary use of nanoparticles as mere vectors for delivery of either the therapeutic drug or the imaging contrast agent, inorganic vectors combine these roles and create a multifunctional system having both these capabilities. In addition, the inherent properties of the core imaging agents, such as iron oxide, gold nanoparticles and quantum dots, allows these nanoparticles to function in alternative anticancer therapies such as hyperthermia,77–79 radiation and photodynamic therapy.80,81

10.3.1 Functionalised magnetic (iron oxide) nanoparticles The concept of using magnetic nanoparticles for drug delivery was coined by Widder, Senyei and colleagues in 1978.82 There are many different methods for synthesising the magnetic nanoparticles. The most commonly used is the precipitation-based approach, either by co-precipitation or reverse micelle synthesis. Fe2+(aq.) + Fe3+(aq.) + 8OH−(aq.) → Fe3O4(s) + 4H2O(I)

[10.1]

Over the past three decades, numerous nanoparticulate iron oxide preparations have been reported and used in cellular therapy, tissue repair, drug delivery, and hyperthermia,83,84 magnetic resonance imaging (MRI)85,86 and, more recently, as sensors for metabolites and other biomolecules.87,88 Recent research has focused on the targeted delivery of iron oxide nanoparticles, which has led to active exploration of the application of magnetic nanoparticles in biomedicine. A brief overview of the functionalisation of magnetic nanoparticles and their applications is shown in Fig. 10.6. Illustrative examples of their applications are shown in Table 10.4. Multifunctional MNPs can be fabricated using two strategies: (a)

Molecular functionalisation using antibodies, ligands or receptors, and make them interact with a biological entity with high affinity. (b) Combine magnetic nanoparticles and other functional nanostructures by sequential growth or coating, and provide multifunction in a single entity. Surface modification using biomolecules improves the target localisation of the MNPs and imparts stability under physiological conditions. Most clinical preparations (Ferridex®, Combidex®, Resovist®, and AMI-228/ ferumoxytrol) have been based upon dextran or similar carbohydrate coatings. Among them, dextran-coated iron oxide nanoparticles play a

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Potential applications

‘Biomolecules’

‘Nanocomponents’

Specific binding, drug delivery

Antibodies

Bacteria detection, protein separation

Ligands or

QDs

or DNAs

Multimodal imaging 5 nm

‘Nanodrugs’

Drug delivery, MRI

receptors 5 nm

Dyes Multimodal imaging

Potential applications

Metal Multimodal imaging, multivalency

Magnetic nanoparticles 5 nm

10.6 Schematic illustration of functionalisation of magnetic nanoparticles (adapted with permission from © 2009 American Chemical Society89).

major role in clinical cancer imaging and increase the accuracy of cancer nodal staging.91 At first, Zhao et al.92 proved a method of targeting iron oxide nanoparticle to anionic phospholipids that present on the surface of apoptotic cells, for visualising the therapeutic efficacy of a cancer treatment. In that study, the C2 domain of synaptotagmin-I was imparted on the surface of the nanoparticle as the apoptotic targeting moiety. Folid acidbased drug targeting was also attempted by creating a folic acid moiety on the surface of MNPs for improving targeting and intracellular uptake of BT 20 breast cancer cells in vitro.93 There are several methods to incorporate drugs into targeted MNPs. Drugs can be linked to the carrier coating, deposited in the surface layer or trapped within the MNPs themselves.94,95 By considering the targeting efficacy of folic acid and its analogue, Kohler et al. showed that MNPs can be multifunctionalised using methotrexate, an anticancer agent.96 Methotrexate is analogous to folic acid and hence methotrexate functionalised MNPs can be internalised by recognising folic acid receptors. Inside the cell, lysosomal pH cleaves methotrexate from the surface and exerts chemotherapeutic effect for cancer irradiation. The multifunctionality of MNPs, combining such tumour targeted imaging with drug delivery, is an important step in the creation of all-in-one cancer therapy. They act as a potential system for the delivery of chemotherapeutic drugs to the tumour, retaining enough magnetic strength for imaging contrast enhancement. To induce properties of MNPs to minimise or eliminate opsonisation, MNPs have been functionalised using amphiphilic polymers, such as poloxamers and other PEG derivatives. Also, functional groups of modified PEG allow for bioconjugation of various ligands or therapeutic agents to MNPs.97 A more recent approach has involved the formulation of MNPs using oleic acid,

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Monoclonal antibody-610 Antibody to carcinoembryonic antigen (CEA) Monoclonal antibody L6 Transferrin

USPIO (Cerdan et al., 1989)

SPIO (Simberg et al., 2007)

Biofunctional PEG-SPIO (Sun et al., 2006) SPIO (Leuschner et al., 2006)

Iron oxide nanocrystals (Fe3O4) (Huh et al., 2005) SPIO (Kohler et al., 2005)

Folic acid

Dextran-coated superparamagnetic maghemite (γ-Fe2O3) nanocrystals (Sonvico et al., 2005) Ferumoxides (SPIO) (Toma et al., 2005)

LHRH receptors

Luteinizing hormone releasing hormone (LHRH) CREKA peptide

Clotted plasma proteins

Folate receptors

Colorectal tumour antigens HER2/neu receptors

Underglycosylated mucin-1 antigens (uMUC-1) Folate receptors

HER2/neu receptors

Transferrin receptors

Surface antigens

CEA

Surface antigens

Target

Folic acid

Methotrexate

Monoclonal antibody A7 Herceptin

Monoclonal antibody HER2/neu EPPT peptide

Streptavidin-conjugated SPIO (Artemov et al., 2003) CLIO-NH2 (Moore et al., 2004)

USPIO (Kresse et al., 1998)

MINO (Remsen et al., 1996)

SPIO (Tiefenauer et al., 1996)

Targeting ligands

Iron oxide nanoparticle

Breast cancer

Human cervical cancer cells Human cervical cancer cells Breast cancer

NIH3T6.7

Colorectal carcinoma

Breast, colon, pancreas and lung cancer cell lines Human epithelial mouth carcinoma

Intracranial tumour LX-1 Rat mammary carcinoma Breast cancer

Colon carcinoma cell lines Colon tumour

Tumour

In vivo

In vivo

In vitro

In vitro

In vivo

In vivo

In vitro

In vivo

In vitro

In vivo

In vivo

In vivo

In vitro

Experimental

Table 10.4 Illustrative examples of functionalised magnetic nanoparticles and their applications (adapted with permission from ©2008 Peng et al.90)

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stabilising with Pluronic F-127 to form a stable, water-dispersible system.98 In this study, Tapan et al. showed that MNPs can be loaded with waterinsoluble anticancer therapeutics (Doxorubicin and paclitaxel) with high efficiency, either alone or in combination, for synergistic activity while retaining their MRI property. Besides combining imaging with therapy, drug-loaded MNPs raise the potential to magnetically guide NPs to deposit drugs at the indented targeted site.99 However in vivo data are not conclusive enough to prove clinical success. There occurs dampening of external magnetic field with increasing depth in the biological environment, which limits magnetic targeting in in vivo conditions. Combination of magnetic nanomaterials and other nanocomponents nanostructures (QD, gold nanoparticle) leads to the creation of a single nanoentity (a dumbbell-like nanoparticle; DBNPs) having multifunctional properties. In this case, different NP surfaces facilitate the controlled functionalisation of each NP with a targeting agent or a therapeutic drug and make them promising multifunctional probes for diagnostic and therapeutic applications. Examples of such systems include Au–Fe3O4, Ag–Fe3O4, Au–FePt and CdSe–Fe2O3 NPs.100 Figure 10.7a illustrates the surface functionalisation of Au–Fe3O4 NPs using epidermal growth factor receptor antibody (EGFRA) via PEG by (1-ethyl-3-(3-dimethylaminopropyl) (a) O

EGFRA–NH–CO–PEG(3000)–CO–NH

–S–PEG(2000)–NH2 O

(b)

(c)

10 nm

10.7 (a) Schematic illustration of surface functionalisation of the Au– Fe3O4 DBNPs. (b) TEM image of the 8–20 nm Au–Fe3O4 NPs after surface modification. (c) Reflection images of the 8–20 nm Au–Fe3O4 labelled A431 cells (adapted with permission from © 2009 Wiley-VCH Verlag GmbH101).

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carbodiimide/N-hydroxysuccinimide (EDC/NHS) chemistry on the Fe3O4 side. Figure 10.7b is the TEM image of the NPs after surface modification. Since epidermal growth factor receptor (EGFR) over-express on the surface of A431, epidermal growth factor receptor antibody (EGFRA) conjugated DBNPs can be used as a potential candidate for early diagnosis of and therapies for numerous cancers, including breast and lung cancers. Additional to the ability of MNPs to deliver drug and image tumour in one multifunctional system, they can also be exploited for guided hyperthermia.102,103 With the ability to localise to a great extent in the tumour region, functionalised MNPs impart high efficacy for local heat conduction. The method involves dispersing MNPs throughout the target tissue and then applying an AC magnetic field of sufficient strength and frequency. Heat is generated around the MNPs as a result of hysterisis loss and spreads immediately into the surrounding diseased tissue; cell apoptosis or necrosis results on increasing the temperature above 42 °C for 30 minutes or more. Local hyperthermia also enhances the perfusion of systematically administered drugs into the core of the solid tumour mass. Most of the in vivo results support the usefulness of this principle of effective tumour shrinkage.

10.3.2 Functionalised gold nanoparticles Gold nanoparticles have recently emerged as an attractive candidate for delivery of various payloads into their targets. The gold nanoparticles (GNPs) exploit their unique physical and chemical properties for transporting and releasing drug molecules. The important characteristics of gold NPs include: (a) Gold core is basically inert and non-toxic,104 (b) monodisperse gold nanoparticles having sizes ranging from 1 nm to 150 nm can be easily synthesised by various methods, (c) multifunctional gold NPs can be created through thiol linkages, and (d) photophysical properties can trigger drug release at remote places.105 GNPs with varying core sizes are prepared by the reduction of gold salts in the presence of appropriate stabilising agents that prevent particle agglomeration. A wide variety of monolayer protected clusters (MPCs) can be used for the functionalisation of GNPs. A schematic representation is shown in the Fig. 10.8. Functionalised gold nanoparticles represent highly attractive and promising candidates in the application of drug delivery, owing to their unique dimensions, tunable functionalities on the surface, and controllable drug release. Combined imaging and therapeutic use of a gold nano shell has been proven in several cancer models, both in vivo and in vitro.107 In this case, silica core nanoparticles are surrounded by a gold coating. Since thiol groups have strong affinity towards gold nanoparticles, thiolated PEG can be easily assembled onto the nanoshell surface and allows for the incorporation of targeting ligand onto the nanoshell system. It imparts a synergic

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HAuCl4

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HS NaBH4

S S S S S

SSS

SSS

S S S S S

HS

S S S S S

SSS

SSS

S S S S S

HS

S S S S S

SSS

SSS

S S S S S

HS

HS HS

10.8 Schematic representation for the functionalisation of gold nanoparticle using various chemical moieties which are anchored with thiol linkers (-SH) (adapted with permission from © 2008 Elsevier106).

effect on targeting efficiency, combined with passive targeting by the EPR effect. Even though it is possible to impart a variety of antibodies onto the surface of GNPs by direct conjugation,108 the use of polymer molecules having free carboxylic acid or amino groups expands the options for targeting ligand molecules that have no sulfhydryl groups for covalent linkages.109 Andres and co-workers have recently demonstrated folic acid conjugated GNPs for target-specific delivery, using the PEG spacer.110 The study showed that particles were effectively internalised by positive KB cells. In another study, methotrexate was conjugated on gold nanoparticles and Wu, Shiau and co-workers111 demonstrated that it inhibited tumour growth in mouse ascites model of Lewis lung carcinoma. Paciotti et al. in their study, used colloidal gold particle conjugate with TNF-α and demonstrated its targeted anticancer therapeutic in MC 38 colon carcinoma in vivo.112 Glutathionemediated (GSH) release used in clinical practice is an alternative, nonenzymatic strategy for the selective intracellular activation of drug molecules.113–115 Gold nanoparticle can be effectively targeted on specific disease site in conjugation with glutathione moiety. Hong et al. demonstrated cellular delivery and GSH-mediated release of a hydrophobic dye (BODIPY), as a model for hydrophobic drugs, using functionalised gold nanoparticles.116 A multifunctional moiety composed of tetra(ethylene glycol)ylated cationic ligands (TTMA) and fluorogenic ligands (HSBDP) facilitated the crossing of cell-membrane barriers. Generally BODIPYconjugated GNPs are nonfluorescent in nature; during conjugation, gold core quenches fluorescence via energy and/or electron transfer processes.117 On triggering the GNPs with GSH, a fluorescence signal originates and leads to simultaneous imaging and therapy. Figure 10.9 shows a GSHmediated payload release via a place-exchange reaction and its fluorescence imaging. In addition to the surface chemistry of GNPs, their physical properties could be exploited for the delivery application, such as local heating when

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(b) GSH 37 ∞C

GNP

NH3+

–

H N

O O

O

HS

O– O N + F NB F

+

HS

(c)

O N H SH

( )9 ( O

)4N

+

GNP

=

= GSH

GSH

GSH-OEt =

= HSBDP (Drug analogue)

=

= TTMA 0 mM

5 mM

20 mM

[GSH-OEt]

10.9 (a) Schematic depiction of the GSH-mediated payload release via place-exchange reaction. (b) Bright field and fluorescence micrographs of human Hep G2 cells after incubation with GNPs for 96 h. (c) Fluorescence images showing dose-dependent release of the payloads. (Adapted with permission from © 2008 Elsevier.106)

irradiated with light. El-Sayed et al. have recently reported the potential use of GNPs in photothermal destruction of tumours.118 In that study, citrate-stabilised GNPs were coated with anti-EGFR (epidermal growth factor receptor) to target HSC3 cancer cells (human oral squamous cell carcinoma) and the study showed that the functionalisation enhanced photothermal efficacy of GNPs by 20 times.

10.4

Conclusion

There are many exciting advantages for the use of functionalised nanoparticles in drug delivery, but this is challenging because of the complexity of the body and the targeting barriers. Functionalised nanoparticles offer unlimited opportunities by focusing on target molecules and the simultaneous design of nanocarriers. Even though different strategies are employed for functionalisation of the nanocarrriers, it is hard to precisely control the number of functional molecule and the coordination of individual properties of various chemical moieties on the surface of the nanocarriers. Studies are ongoing to develop better strategies for producing multifunctional nanocarriers with uniform surface properties and reproducible functionalisation.

10.5

Acknowledgements

The authors wish to thank the Director of SCTIMST and the Head of the Biomedical Technology Wing, SCTIMST for providing facilities. One of the

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authors (S. Manju) also thanks the Department of Biotechnology, New Delhi, India for financial support.

10.6

References

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48 t.p. thomas, a.k. patri, a. myc, m.t. myaing, j.y. ye, t.b. norris, j.r. baker jr. In vitro targeting of synthesized antibody conjugated dendrimer nanoparticles. Biomacromolecules, 2004; 5:2269–2274. 49 t.m. fahmy, p.m. fong, j. park, t. constable, w.m. saltzman. Nanosystems for simultaneous imaging and drug delivery to T cells. AAPS J., 2007; 9:E171–E180. 50 b.a. khaw, a.l. klibanov, s.m. o’donnell, t. saito, n. nossiff, m.a. slinkin, j.b. newell, h.w. strauss, v.p. torchilin. Gamma imaging with negatively chargemodified monoclonal antibody: Modification with synthetic polymers. J. Nucl. Med., 1991; 32:1742–1752 51 o.c. farokhzad, s. jon, a. khademhosseini, t.t. tran, d.a. lavan, r. langer. Nanoparticle–aptamer bioconjugates: a new approach for targeting prostate cancer cells. Cancer Res., 2004; 64:7668–7672. 52 o.c. farokhzad, j.m. karp, r. langer. Nanoparticle–aptamer bioconjugates for cancer targeting. Expert Opin. Drug Deliv., 2006; 3:311–324. 53 s. ganta, h. devalapally, a. shahiwala, m. amiji. A review of stimuli-responsive nanocarriers for drug and gene delivery. J. Control. Release, 2008; 126: 187–204. 54 s. ni, s.m. stephenson, r.j. lee. Folate receptor targeted delivery of liposomal daunorubicin into tumor cells. Anticancer Res., 2002; 22:2131–2135. 55 x.o. pan, h. wang, r.j. lee. Antitumor activity of folate receptor-targeted liposomal doxorubicin in a KB oral carcinoma murine xenograft model. Pharm. Res., 2003; 20:417–422. 56 s.m. stephenson, w. yang, p.j. stevens, w. tjarks, r.f. barth, r.j. lee. Folate receptor-targeted liposomes as possible delivery vehicles for boron neutron capture therapy. Anticancer Res., 2003; 23:3341–3345. 57 x.q. pan, h. wang, s. shukla, m. sekido, d.m. adams, w. tjarks, r.f. barth, r.j. lee. Boron-containing folate receptor-targeted liposomes as potential delivery agents for neutron capture therapy. Bioconjugate Chem., 2002; 13:435–442. 58 b. stella, s. arpicco, m.t. peracchia, d. desmaele, j. hoebeke, m. renoir, j. d’angelo, l. cattel, p. couvreur. Design of folic acid-conjugated nanoparticles for drug targeting. J. Pharm. Sci., 2000; 89:1452–1464. 59 e.k. park, s.b. lee, y.m. lee. Preparation and characterization of methoxy poly(ethylene glycol)/poly(epsilon-caprolactone) amphiphilic block copolymeric nanospheres for tumor specific folate-mediated targeting of anticancer drugs. Biomaterials, 2005; 26:1053–1061. 60 h. hatakeyama, h. akita, k. maruyama, t. suhara, h. harashima. Factors governing the in vivo tissue uptake of transferrin-coupled polyethylene glycol liposomes in vivo. Int. J. Pharm., 2004; 281:25–33. 61 o. ishida, k. maruyama, h. tanahashi, m. iwatsuru, k. sasaki, m. eriguchi, h. yanagie. Liposomes bearing polyethylene glycol-coupled transferrin with intracellular targeting property to the solid tumors in vivo. Pharm. Res., 2001; 18:1042–1048. 62 j.l. wike-hooley, j. haveman, h.s. reinhold. The relevance of tumor pH to the treatment of malignant disease. Radiother. Oncol., 1984; 2:343–366. 63 p. vaupel, f. kallinowski, p. okunieff. Blood flow, oxygen and nutrient supply, and metabolic microenvironment of human tumors: A review. Cancer Res., 1989; 23:6449–6465.

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64 l.e. gerweck, k. seetharaman. Cellular pH gradient in tumor versus normal tissue: Potential exploitation for the treatment of cancer. Cancer Res., 1996; 56:1194–1198. 65 h. devalapally, d. shenoy, s. little, r. langer, m. amiji. Poly(ethylene oxide)modified poly(beta-amino ester) nanoparticles as a pH-sensitive system for tumor-targeted delivery of hydrophobic drugs: Part 3. Therapeutic efficacy and safety studies in ovarian cancer xenograft model. Cancer Chemother. Pharmacol., 2007; 59:477–484. 66 m.w. dewhirst, z. vujaskovic, e. jones, d. thrall. Re-setting the biologic rationale for thermal therapy. Int. J. Hyperth., 2005; 21:779–790. 67 a.k. gupta, r.r. naregalkar, v.d. vaidya, m. gupta. Recent advances on surface engineering of magnetic iron oxide nanoparticles and their biomedical applications. Nanomed., 2007; 2:23–39. 68 h. feil, y.h. bae, j. feijen, s.w. kim. Effect of co-monomer hydrophilicity and ionization on the lower critical solution temperature of N-isopropylacrylamide copolymers. Macromolecules, 1993; 26:2496–2500. 69 g. molineux. PEGylation: Engineering improved pharmaceuticals for enhanced therapy. Cancer Treat. Rev., 2002; 28:13–16. 70 j.e. chung, m. yokoyama, t. okano. Inner core segment design for drug delivery control of thermo-responsive polymeric micelles. J. Control. Release, 2000; 65:93–103. 71 x.m. liu, y.y. yang, k.w. leong. Thermally responsive polymeric micellar nanoparticles self-assembled from cholesteryl end-capped random poly (Nisopropylacrylamide-co-N,N-dimethylacrylamide): Synthesis, temperaturesensitivity, and morphologies. J. Colloid Interface Sci., 2003; 266:295–303. 72 w.s. shim, j.h. kim, k. kim, y.s. kim, r.w. park, i.s. kim, i.c. kwon, d.s. lee. pH- and temperature-sensitive, injectable, biodegradable block copolymer hydrogels as carriers for paclitaxel. Int. J. Pharm., 2007; 331:11–18. 73 k. kono, k. yoshino, t. takagishi. Effect of poly(ethylene glycol) grafts on temperature-sensitivity of thermo sensitive polymer-modified liposomes. J. Control. Release, 2002; 80:321–332. 74 h.d. han, b.c. shin, h.s. choi. Doxorubicin-encapsulated thermosensitive liposomes modified with poly(N-isopropylacrylamide-co-acrylamide): Drug release behavior and stability in the presence of serum. Eur. J. Pharm. Biopharm., 2006; 62:110–116. 75 j. kim, y. piao, t. hyeon. Multifunctional nanostructured materials for multimodal imaging, and simultaneous imaging and therapy. Chem Soc Rev., 2009; 38:372–390. 76 m. liong, j. lu, m. kovochich, t. xia, s.g. ruehm, a.e. nel, f. tamanoi, j.i. zink. Multifunctional inorganic nanoparticles for imaging, targeting, and drug delivery. ACS Nano., 2008; 2:889–896. 77 q.a. pankhurst, j. connolly, s.k. jones, j. dobson. Applications of magnetic nanoparticles in biomedicine. J. Phys. D: Appl. Phys., 2003; 36:R167– R181. 78 s. mornet, s. vasseur, f. grasset, e. duguet. Magnetic nanoparticle design for medical diagnosis and therapy. J. Mater. Chem., 2004; 14:2161–2175. 79 r.k. visaria, r.j. griffin, b.w. williams, e.s. ebbini, g.f. paciotti, c.w. song, j.c. bischof. Enhancement of tumor thermal therapy using gold nanoparticle-

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assisted tumor necrosis factor-A delivery. Mol Cancer Ther, 2006; 5:1014– 1020. s. brown, e. brown, i. walker. The present and future role of photodynamic therapy in cancer treatment. Lancet Oncol., 2004; 5:497–508. c.h. sibata, v.c. colussi, n.l. oleinick, t.j. kinsella. Photodynamic therapy in oncology. Expert Opin. Pharmacother., 2001; 2:917–927. a. senyei, k. widder, c. czerlinski. Magnetic guidance of drug carrying microspheres. J. Appl. Phys., 1978; 49:3578–3583. c.c. berry, a.s.g. curtis. Functionalisation of magnetic nanoparticles for applications in biomedicine. J. Phys. D: Appl. Phys., 2003; 36:R198–R206 . a. ito, m. shinkai, h. honda, t. kobayashi. Medical application of functionalized magnetic nanoparticles. J. Biosci. Bioengng., 2005; 100:1–11. r. weissleder, a. moore, u. mahmood, r. bhorade, h. benveniste, e.a. chiocca, j.p. basilion. In vivo magnetic resonance imaging of transgene expression. Nature Med., 2000; 6:351–354. h. choi, s. choi, r. zhou, h. kung, i. chen. Iron oxide nanoparticles as magnetic resonance contrast agent for tumor imaging via folate receptor-targeted delivery. Academic Radiol., 2004; 11:996–1004. j. grimm, j.m. perez, l. josephson, r. weissleder. Novel nanosensors for rapid analysis of telomerase activity. Cancer Res., 2004; 64:639–643. g.y. kim, l. josephson, r. langer, m.j. cima. Magnetic relaxation switch detection of human chorionic gonadotrophin. Bioconjug. Chem., 2007; 18: 2024–2028. j. gao, h. gu, b. xu. Multifunctional magnetic nanoparticles: design, synthesis, and biomedical applications. Accounts Chem. Res., 2009; 42:1097–1107. x.h. peng, x. qian, h. mao, a.y. wang, z. chen, s. nie, d.m. shin. Targeted magnetic iron oxide nanoparticles for tumor imaging and therapy. Int. J. Nanomed., 2008; 3:311–321. m. ferrari. Cancer nanotechnology: Opportunities and challenges. Nat. Rev. Cancer, 2005; 5:161–171. m. zhao, d.a. beauregard, l. loizou, b. davletov, k.m. brindle. Non-invasive detection of apoptosis using magnetic resonance imaging and a targeted contrast agent. Nat. Med., 2001; 7:1241–1244. y. zhang, j. zhang. Surface modification of monodisperse magnetite nanoparticles for improved intracellular uptake to breast cancer cells. J. Coll. Inter. Sci., 2005; 283:352–357. r. jurgons, c. seliger, a. hilpert, l. trahms, s. odenbach, c. alexiou. Drug loaded magnetic nanoparticles for cancer therapy. J. Phys.: Condens. Matter, 2006; 18:S2893–S2902. s. purushotham, p.e.j. chang, h. rumpel, i.h.c. kee, r.t.h. ng, p.k.h. chow, c.k. tan, r.v. ramanujann. Thermoresponsive core–shell magnetic nanoparticles for combined modalities of cancer therapy. Nanotechnology, 2009; 20:305101. k. nathan, s. conroy, w. jassy, m. zang. Methatrexate modified superparamagnetic nanoparticle and their cellular uptake into human cancer cells. Langmuir, 2005; 21:8858–8864. c. sun, r. sze, m. zhang. Folic acid–PEG conjugated superparamagnetic nanoparticles for targeted cellular uptake and detection by MRI. J. Biomed. Mater. Res. Part A, 2006; 78A: 3550–3557.

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98 k.j. tapan, j. richey, m. strand, d.l. leslie-pelecky, c.a. flask, v. labhasetwar. Magnetic nanoparticles with dual functional properties: Drug delivery and magnetic resonance imaging. Biomaterials, 2008; 29:4012–4021. 99 r. asmatulu, m.a. zalich, r.o. claus, j.s. riffle. Synthesis, characterization and targeting of biodegradable magnetic nanocomposite particles by external magnetic fields. J. Magnetism Magnetic Mater., 2005; 292:108–119. 100 s.t. selvan, p.k. patra, c.y. ang, j.y. ying. Synthesis of silica-coated semiconductor and magnetic quantum dots and their use in the imaging of live cells. Angew. Chem. Int. Ed., 2007; 46:2448–2452. 101 c. wang, c. xu, h. zeng, s. sun. Recent progress in syntheses and applications of dumbbell-like nanoparticles. Adv. Mater., 2009; 21:3045–3052. 102 q.a. pankhurst, j. connolly, s.k. jones, j. dobson. Applications of magnetic nanoparticles in biomedicine. J. Phys. D, Appl. Phys., 2003; 36:R167– R181. 103 f. sonvico, s. mornet, s. vasseur, c. dubernet, d. jaillard, j degrouard, j. hoebeke, e. duguet, p. colombo, p. couvreur. Folate-conjugated iron oxide nanoparticles for solid tumor targeting as potential specific magnetic hyperthermia mediators: Synthesis, physicochemical characterization, and in vitro experiments. Bioconjug. Chem., 2005; 16:1181–1188. 104 e.e. connor, j. mwamuka, a. gole, c.j. murphy, m.d. wyatt. Gold nanoparticles are taken up by human cells but do not cause acute cytotoxicity. Small, 2005; 1:325–327. 105 a.g. skirtach, a.m. javier, o. kreft, k. kohler, a.p. alberola, h. mohwald, w.j. parak, g.b. sukhorukov. Laser-induced release of encapsulated materials inside living cells. Angew. Chem. Int. Ed., 2006; 45:4612–4617. 106 p. ghosh, g. han, m. de, c.k. kim, v.m. rotello. Gold nanoparticles in delivery applications. Adv. Drug Deliv. Rev., 2008; 60:1307–1315. 107 a.m. gobin, m.h. lee, n.j. halas, w.d. james, r.a. drezek, j.l. west. Near-infrared resonant nanoshells for combined optical imaging and photothermal cancer therapy. Nano. Lett., 2007; 7:1929–1934. 108 a.j.d. pasqua, r.e. mishler ii, y. ship, j.c. dabrowiak, t. asefa. Preparation of antibody-conjugated gold nanoparticles. Mater. Lett., 2009; 63:1876–1879. 109 w. liu, x. yang, w. huang. Catalytic properties of carboxylic acid functionalizedpolymer microsphere-stabilized gold metallic colloids. J. Colloid Interface Sci., 2006; 304:160–165. 110 v. dixit, j. van den bossche, d.m. sherman, d.h. thompson, r.p. andres. Synthesis and grafting of thioctic acid–PEG–folate conjugates onto Au nanoparticles for selective targeting of folate receptor-positive tumor cells. Bioconjug. Chem., 2006; 17:603–609. 111 y.h. chen, c.y. tsai, p.y. huang, m.y. chang, p.c. cheng, c.h. chou, d.h. chen, c.r. wang, a.l. shiau, c.l. wu. Methotrexate conjugated to gold nanoparticles inhibits tumor growth in a syngeneic lung tumor model. Mol. Pharm., 2007; 4:713–722. 112 g.f. paciotti, l. myer, d. weinreich, d. goia, n. pavel, r.e. mclaughlin, l. tamarkin. Colloidal gold: a novel nanoparticle vector for tumor directed drug delivery. Drug Deliv., 2004; 11:169–183. 113 h. sies. Glutathione and its role in cellular functions. Free Radic. Biol. Med., 1999; 27:916–921.

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114 d.p. jones, j.l. carlson, v.c. mody, j.y. cai, m.j. lynn, p. sternberg. Redox state of glutathione in human plasma. Free Radic. Biol. Med., 2000; 28:625–635. 115 r. hong, g. han, j.m. fernandez, b.j. kim, n.s. forbes, v.m. rotello. Glutathione mediated delivery and release using monolayer protected nanoparticle carriers. J. Am. Chem. Soc., 2006; 128:1078–1079. 116 k.e. sapsford, l. berti, i.l. medintz. Materials for fluorescence resonance energy transfer analysis: Beyond traditional donor–acceptor combinations. Angew. Chem. Int. Ed., 2006; 45:4562–4588. 117 e. dulkeith, a.c. morteani, t. niedereichholz, t.a. klar, j. feldmann, s.a. levi, f.c.j.m. van veggel, d.n. reinhoudt, m. moller, d.i. gittins. Fluorescence quenching of dye molecules near gold nanoparticles: Radiative and nonradiative effects. Phys. Rev. Lett., 2002; 89:203002. 118 x. huang, w. qian, i.h. el-sayed, m.a. el-sayed. The potential use of the enhanced nonlinear properties of gold nanospheres in photothermal cancer therapy. Laser Surg. Med., 2007; 39:747–753.

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11 Biocompatibility of materials and its relevance to drug delivery and tissue engineering T. C H A N DY, 3M Drug Delivery Systems, USA

Abstract: All materials intended for application in humans as biomaterials, medical devices, or prostheses undergo tissue responses when implanted into living tissue. Similarly when blood contacts a biomaterial surface, a variety of blood components interfere with the surface, leading to thrombosis or complement activation. This review first describes fundamental aspects of tissue/blood responses to materials, which are commonly described as the tissue/blood response continuum. These actions involve fundamental aspects of tissue responses, including injury, inflammatory and wound healing responses, foreign body reactions, and fibrous encapsulation of the biomaterial, medical device or prosthesis. The second part of this review describes the biocompatibility of materials being used in medical devices and prostheses to suit their applications. The review includes an emphasis on the biocompatibility of biomaterials being used in drug delivery and tissue engineering applications. It also summarizes the use of scaffolds in the dual role of structural support for cell growth and vehicle for controlled release of tissue inductive factors, or DNA encoding for these factors. The confluence of molecular and cell biology, materials science and engineering provides the tools to create controllable microenvironments that mimic natural developmental processes and direct tissue formation for experimental and therapeutic applications. The chapter ends with recent approaches towards combination therapy devices, such as stent modifications with surface engineering and site-specific drug delivery. Key words: biocompatibility, inflammatory response, surface modifications, biostability, drug delivery, growth hormones, polymer scaffold, tissue engineering, combination devices.

11.1

Biocompatibility of materials and medical applications

The biocompatibility of a long-term implantable medical device refers to the ability of the device to perform its intended function, with the desired degree of incorporation in the host, without eliciting any undesirable local or synthetic effects. Practically speaking, the evaluation of biological response to a medical device is carried out to determine that the medical device performs as intended and presents no significant harm to the patient 301 © Woodhead Publishing Limited, 2010

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or user. Thus the goal of biological response evaluation is to predict whether a biomaterial, medical device, or prostheses presents potential harm to the patient or user, by evaluation under conditions that simulate clinical use. In almost all situations, the practical consequence is that we select devices that irritate the host the least, through choice of the most inert and least toxic materials and the most appropriate mechanical design.

11.1.1 Fundamental aspects of tissue response to materials The process of implantation of a biomaterial, prosthesis, or medical device results in injury to surrounding tissues or organs.1,2 The response to injury is dependent on multiple factors, including the extent of the injury, the loss of basement membrane structures, blood–material interactions, provisional matrix formation, the extent or degree of cellular necrosis, and the extent of the inflammatory response. These events in turn may occur very early, i.e. within two to three weeks from the time of implantation, and may affect the subsequent perturbation of homeostatic mechanisms that lead to the cellular cascades of wound healing and the subsequent capsule development. These events are summarized in Table 11.1. The end-stage healing response to biomaterials is generally fibrosis or fibrous encapsulation. A thin fibrous capsule is inevitable for the longterm stability of the implant and a thick capsule shows the irritation of the tissue due to the implant. Local systemic factors may play a role in the wound-healing response to biomaterials or implants. Local factors include the site (tissue or organ) of implantation, the adequacy of blood supply, and the potential for infection. Systemic factors may include nutrition, hematological and immunological derangements, glucocortical steroids, and pre-existing diseases such as atherosclerosis, diabetes and infection.3,4

Table 11.1 Sequence of host reactions following implantation of medical devices Injury Blood–material interactions Provisional matrix formation Acute inflammation Granulation tissue Foreign body reaction Fibrosis/fibrous capsule development

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11.1.2 Blood–material interactions and initiation of the inflammatory response The normal host response to an implant includes trauma, inflammation, the immune system’s reaction, and eventual healing or scarring. Biomaterials exhibiting a lack of biocompatibility can induce many complications, which might include long-lasting chronic inflammation or cytotoxic chemical buildup. Blood–material interactions and the inflammatory response are intimately linked and, in fact, early responses to injury involve mainly blood and the vasculature.3–6 The host response to biomaterials or to an implant is demonstrated in Table 11.2. Regardless of the tissue or organ into which a biomaterial is implanted, the initial inflammatory responses, thrombi and/ or blood clots, are formed due to the injury to vascular connective tissue (Table 11.2). Thrombus formation involves activation of extrinsic and intrinsic coagulation systems, the complement systems, the fibrinolytic system, the kiningenerating system, and platelets.6,7 The adhesion of circulating platelets at sites of injury on the vessel wall where subendothelium, in particular collagen, is exposed to the flowing blood, is an important step in the formation of a hemostatic plug. Activation of platelets by contact with biomaterials is a key event in the thromboembolic complications of prosthetic devices in contact with blood. It is known that a film of plasma protein adsorbs on biomaterials exposed to blood and that this event proceeds interaction of the surface with blood cells.8,9 Platelet adhesion to biomaterial surfaces occurs in various steps, including initial attachment, spreading, release of granule contents, and platelet aggregation.

Table 11.2 Host response characteristics Protein adsorption and desorption characteristics Complement activation Platelet adhesion, activation and aggregation Activation of intrinsic clotting cascade Neutrophil activation Fibroblast behavior and fibrosis Microvascular changes Macrophage activation, foreign body giant cell production Osteoblast/osteoclast responses Endothelial proliferation Antibody production, lymphocyte behavior Acute hypersensitivity/anaphylaxis Delayed hypersensitivity Genotoxicity, reproductive toxicity Tumor formation

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Vroman et al.10 extensively studied the initial phase of blood–biomaterial interactions. They postulated that on hydrophilic surfaces (e.g. glass), fibrinogen is deposited within seconds, along with traces of high molecular weight kininogen (HMWK) and Factor XII. Then more HMWK and Factor XII arrive and displace fibrinogen. Platelets adhere most where fibrinogen remains. Thus, fibrinogen plays a key role in platelet–biomaterial attachment and subsequent thrombus formation. An early feature of activation of platelets by soluble agonists, such as adenosine diphosphate (ADP) or thrombin, is exposure of specific membrane glycoproteins (GPIIb-IIIa), to which fibrinogen molecules bind with high affinity.11,12 It should be noted that fibrinogen dissolved in plasma does not induce platelet aggregation but rather acts as a cofactor in the process, once the platelet GPIIb-IIIa receptors are activated. Fibrinogen normally circulates in peaceful coexistence with unactivated platelets without any obvious interaction.12,13

11.1.3 Surface modifications to improve biocompatibility of materials Surface modification of biomaterials is one approach for improving their functionalities for best performance. Various investigators14–17 have used physical, chemical or biological methods to generate an inert or passive interface with blood/tissue reactions. These include albumin, collagen, laminin, fibronectin, heparin coatings, hirudin, antiplatelet drugs, prostaglandins and growth factors, for normal healing and naturalization.15,16 Table 11.3 gives a summary of the surface modifications used for improving biocompatibility of implants. However, successful results have not yet been reported for small-caliber artificial vascular grafts. It seems that modifications of these prostheses surfaces which stimulate endo-thelialization could reduce thrombosis, eliminate platelet deposition, resist bacterial infection and extend graft patency.18,19

Table 11.3 Surface modifications for improving biocompatibility Physical treatments

Plasma glow, radiation, UV

Chemical modifications

Hydrogels (PHEMA, PEG, hyaluronic acid, synthetic phosphoryl choline derivatives, silicone coatings, etc.)

Biological modifications

Albumin coatings, gelatin, heparin, LMW heparin, hirudin, collagen IV, laminin-5, growth factors, PC, etc. Tissue engineering

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Table 11.4 Effect of immobilized biomolecules on the adhesion of platelets and fibrinogen absorption on PTFE and Dacron surfaces

Surfaces

Platelet adhesion 30 min (mm2 ± SD)

PTFE (untreated) CL-PTFE-PGE1 CL-PTFE-heparin CL-PTFE-PC Dacron (untreated) CL-Dacron-PGE1 CL-Dacron-heparin CL-Dacron-PC

62.92 21.65 15.00 16.58 48.93 14.05 17.50 13.58

± ± ± ± ± ± ± ±

4.9 3.3 3.5 4.2 5.7 3.5 3.3 3.2

Fibrinogen adsorption (μg/cm2) 1.08 0.41 0.26 0.21 0.86 0.27 0.22 0.19

± ± ± ± ± ± ± ±

0.21 0.02 0.01 0.02 0.04 0.01 0.02 0.01

CL, collagen and laminin-immobilized surface. Compiled from reference 20.

Enhancement of antithrombogenicity at the luminal surface and tissue regeneration at the outer surface may lead to a vital and functional artificial graft. Surface modifications that have been used to enhance endothelial cell seeding on vascular prostheses include immobilization of fibronectin, laminin, collagen, growth factors and peptides.18–20 An ideal coating process should have a normal healing process and naturalization. It seems that an ideal healing process is as follows: as the luminal surface coating is biodegraded, while it maintains its thrombus-free character, tissue formation, including endothelial regeneration progresses. When this coating is completely biodegraded, generated tissues are expected to replace it. Doi and Matsuda21 prepared microporous polyurethane grafts that were then coated with photoreactive gelatin, fibroblast growth factor and heparin, and were photocured by UV irradiation. Histological evaluation of these grafts showed that the neoarterial wall of the fibroblast growth factor/ heparin-impregnated graft, was much thicker than that of the noncoated graft and was formed with endothelial cells. Thus, the co-immobilization of growth factor and heparin significantly accelerated neoarterial regeneration and patency rates of small-caliber grafts. Chandy et al.20 prepared a series of surface coatings by modifying the argon plasma-treated PTFE and Dacron grafts with Collagen IV and laminin, and subsequently immobilizing bioactive molecules such as PGE1, heparin or phosphatidyl choline via the carbodiimide functionalities. Table 11.4 provides information on platelet adhesion and fibrinogen adsorption to various modified PTFE and Dacron surfaces.20 The in vitro studies showed that the fibrinogen adsorption and platelet adhesion on modified grafts were significantly reduced. This work proposed that the surface grafting of matrix components (Collagen IV and laminin) and subsequent immobilization of bioactive molecules (PGE1,

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Biointegration of medical implant materials Table 11.5 Biomaterials and components relevant to in vivo assessment of tissue compatibility The material(s) of manufacture Intended additives, process contaminants and residues Leachable substances Degradation products Other components and their interactions in the final product The properties and characteristics of the final product

heparin or phosphatidyl choline) changed the surface conditioning of vascular grafts and subsequently improved their biocompatibility.

11.1.4 Biostability and biocompatibility of polymeric materials In the selection of biomaterials to be used in device design and manufacture, the first consideration should be fitness for purpose with regard to characteristics and properties of the biomaterials, which include chemical, toxicological, physical, electrical, morphological and mechanical properties.6 Table 11.5 presents a list of biomaterial components and characteristics that may impact upon the overall biological responses of the medical device. Therefore, knowledge of these components in the medical device, i.e. final product, the duration of the exposure of the device to the tissues and the degradation products, is necessary. The range of potential biological hazards is broad and may include short-term effects, long-term effects, or specific toxic effects, which should be considered for every material and medical device.22 Polyurethanes are used as biomaterials for a variety of applications22,23 such as pacemaker lead insulators, catheters, total artificial hearts, and heart valves. The popularity of polyurethanes for biomedical applications stems from their excellent physical properties and good biocompatibility. Biomaterials, in order to be successful as implant devices, should be well accepted by the host system as well as not exert any adverse effect on the host. However, several studies have demonstrated that the polyether soft segment is susceptible to oxidation after extended periods in vivo.23–25 It is now generally accepted that the reactive oxygen intermediates released by adherent macrophages and foreign body giant cells initiate the observed biodegradation. Because of the known susceptibility of poly(ether urethanes) (PEU) to oxidation, poly(carbonate urethanes) (PCU) are currently being examined as more biostable replacements for long-term implants.25,26 Several in vitro and early in vivo studies have shown that polyurethanes with polycarbonate soft segments are more biostable than

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comparable PEUs. Thus it is necessary to select the biomaterial(s), knowing the projected life of the implant, i.e. for short-term or long-term use.

11.2

Biomaterials for controlled drug delivery

Controlled drug delivery occurs when a polymer, whether natural or synthetic, is judiciously combined with a drug or other active agent in such a way that the active agent is released from the material in a predesigned manner. The release of the active agent may be constant over a long period, it may be cyclic over a long period, or it may be triggered by the environment or other external events. In any case, the purpose behind controlling the drug delivery is to achieve more effective therapies while eliminating the potential for both under- and overdosing.27 Other advantages of using controlled-delivery systems can include the maintenance of drug levels within a desired range, the need for fewer administrations, optimal use of the drug in question, and increased patient compliance. While these advantages can be significant, the potential disadvantages cannot be ignored: the possible toxicity or nonbiocompatibility of the materials used, undesirable by-products of degradation, any surgery required to implant or remove the system, the chance of patient discomfort from the delivery device, and the higher cost of controlled-release systems compared with traditional pharmaceutical formulations. In recent years, controlled drug delivery formulations and the polymers used in these systems have become much more sophisticated, with the ability to do more than simply extend the effective release period for a particular drug. For example, current controlled-release systems can respond to changes in the biological environment and deliver – or cease to deliver – drugs based on these changes. In addition, materials have been developed that should lead to targeted delivery systems, in which a particular formulation can be directed to the specific cell, tissue, or site where the drug it contains is to be delivered.27,28 While much of this work is still in its early stages, emerging technologies offer possibilities that scientists have only begun to explore.

11.2.1 Polymers used in drug delivery A range of materials has been employed to control the release of drugs and other active agents. To be successfully used in controlled drug delivery formulations, a material must be chemically inert and free of leachable impurities. It must also have an appropriate physical structure, with minimal undesired aging, and be readily processable. Some of the materials that are currently being used or studied for controlled drug delivery are shown in Table 11.6. However, in recent years, additional polymers designed primar-

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Biointegration of medical implant materials Table 11.6 Polymers used for drug delivery applications Polyurethanes Poly(siloxanes) or silicones Poly(2-hydroxy ethyl methacrylate) Poly(N-vinyl pyrrolidone) Poly(methyl methacrylate) Poly(vinyl alcohol) Poly(acrylic acid) Polyacrylamide Poly(ethylene-co-vinyl acetate) Poly(ethylene glycol) Poly(methacrylic acid)

Table 11.7 Degradable polymers used for drug delivery Polylactides (PLA) Polyglycolides (PGA) Poly(lactide-co-glycolides) (PLGA) Polyanhydrides Polyorthoesters Chitosan

ily for medical applications have entered the arena of controlled release. Many of these materials are designed to degrade within the body, and are shown in Table 11.7. The greatest advantage of these degradable polymers is that they are broken down into biologically acceptable molecules that are metabolized and removed from the body via normal metabolic pathways.29,30 However, biodegradable materials do produce degradation byproducts that must be tolerated with little or no adverse reactions within the biological environment.

11.2.2 Modified polymers for drug delivery Recurrent luminal narrowing (restenosis) as a result of excessive intimal hyperplasia remains a major limiting factor for the long-term success of vascular interventions.31,32 It has been shown that to inhibit vascular smooth muscle cell proliferation, drugs must be used at a high concentration for a prolonged period of time. Because of the focal nature of restenosis, local delivery systems, such as polymer-based perivascular techniques and catheter-based cardiovascular techniques, have been employed.32–35 Various geometrical devices, such as polymeric matrices, microspheres, and circumferential wraps have been developed for perivascular local delivery.34

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Poly(lactic acid) (PLA) and copolymers of lactic and glycolic acids are wellknown biodegradable and histocompatible aliphatic polymers. They are commonly used as biodegradable sutures, and they have more recently contributed to the reconstruction of deficient or injured organs and to improving galenic formulations.35,36 During the past few years, several techniques for drug encapsulation have been developed that use aliphatic polyesters. Studies of Bazile et al.37 indicated that hydrophilic coatings with poly(ethylene glycol) (PEG) can increase the blood half-life of PLA nanospheres in rats up to several hours. The main advantage of surface-coated microspheres, in comparison with other long-circulating systems, is their shelf-stability and their ability to control the release of the encapsulated drug compound.38,39 Biodegradable microspheres containing taxol were formulated with PLA–PEG polymers as a sustained drug delivery system for the control of restenosis.40 The release of the drug was sustained over four weeks in vitro. These studies suggest that biocompatible PEG-coated PLA microspheres provide a near-zero-order, in vitro release of taxol for therapeutic applications.40 The amount of taxol release from PLA–PEG microspheres was 14.4 mg/mg within 30 days of dissolution compared with 10.7 mg/mg for PLA–PVAlc.40 In other words, PEG-coated microspheres release more taxol to the site for better therapeutics.

11.2.3 Polymer co-matrix system for combination drug delivery Controlled release of appropriate drugs alone and in combinations is one approach for treating coronary obstructions, balloon angioplasty, restenosis associated with thrombosis, and calcification. Chandy et al.41 have demonstrated the possibility of encapsulating taxol-loaded polylactic acid (PLA) microspheres within heparin–chitosan spheres to develop a prolonged release co-matrix form. The co-matrix system was fabricated from taxolloaded PLA microbeads encapsulated in chitosan beads. A controlled delivery of taxol/heparin was achieved by coating the PLA–chitosan co-matrix with polyethylene glycol (Table 11.8). That study41 also highlighted the use of a co-matrix system for targeting system drug combinations with least side effects for the therapeutic applications.

11.2.4 Biocompatible coatings for bioactive protein delivery Proteins and enzymes represent a growing and promising field of therapeutics and are currently administered by injection. Protein delivery from biodegradable polymer systems has been a challenging area of research because of the necessity of improving the delivery of newly developed

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Biointegration of medical implant materials Table 11.8 Amount of taxol and heparin released from PEG-coated PLA–chitosan co-matrix system

Time (days)

Amount of taxol (μg/mg bead)

1 5 15 30 40 60

0.83 1.06 2.17 2.71 3.82 4.08

± ± ± ± ± ±

0.03 0.12 0.08 0.02 0.01 0.13

Amount of heparin (units/mg bead) 13.1 17.7 23.6 25.8 26.0 27.1

± ± ± ± ± ±

0.37 0.62 0.05 1.34 1.4 0.86

macromolecular drugs and antigens.42 During the last few years, several techniques for drug encapsulation have been developed that currently use aliphatic polyesters. As previously mentioned, Bazile et al.37 found that hydrophilic coatings with PEG could increase the blood half-life of PLA nanospheres in rats up to several hours. Chandy et al.43 encapsulated human serum albumin (HSA) and thrombin (Thr) in poly(ethylene glycol) (PEG)coated, monodisperse, biodegradable microspheres with a mean diameter of about 10 micrometer. The PLA–PEG microspheres demonstrated an initial burst release followed by a constant slow release of HSA and Thr for a period of 20 days.43 The main advantage of PEG-coated nanospheres over other longcirculating systems is their shelf-stability and ability to control the release of an encapsulated compound.36,37,42,43 PEG appears to work as a protective colloid for the emulsion droplets during the preparation. The PEG molecules adsorbed on the surface of the droplets prevented coalescence of the droplets. Therefore, it appears that the PEG coating can increase the payload of drugs and ensure better stabilization. Nagaoka et al.44 synthesized a graft copolymer of methacrylates with PEG and found the resulting polymer to be quite nonthrombogenic. PEG-grafted polymer surfaces have also been shown to reduce protein adsorption and are highly resistant to mammalian and bacterial cell adhesion.36,39 Therefore, these new microcapsules fabricated from PLA/PEG may serve to provide controlled protein delivery and immunoprotection, whereas the outer layer of PEG may serve to enhance biocompatibility and reduce biodegradation. PEGcoated PLA microspheres show great potential for protein-based drug delivery. Peptide drug delivery by routes other than the parenteral one has gained much attention in recent years. Many studies are currently being conducted to address other absorption sites for peptide delivery through different targeting mechanisms, such as entrance via the Pyers patch or

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mucoadhesion.45,46 Oral delivery is the easiest method of administration, and allows for a more varied load to be released; however, proteins are quickly denatured and degraded in the hostile environment of the stomach. A potential solution to these problems is the use of microencapsulation processes for oral release of therapeutic agents.46 The protein is encapsulated in a core material that is covered by a biocompatible, semipermeable membrane. The membrane controls the diffuse release rate of the protein from the capsule to the surrounding medium while protecting the remaining encapsulated protein from biodegradation. Alginates, chitosan and polyethylene glycol matrices have been reported potentially useful for medical and pharmaceutical applications such as artificial skin, artificial kidney, cell encapsulation, and as a drug carrier for target delivery.43–46 A polyethylene glycol-coated chitosan/calcium alginate system was used to encapsulate albumin and hirudin for use as an oral delivery system.45 The protein release was less in stomach pH (acidic), and these acid-treated capsules had released almost all the entrapped proteins into intestinal media (pH 7.4) within 6 hours.45 The released hirudin showed their biological activity when tested with specific coagulation assays.45 This study proposed that chitosan–alginate microbeads may be used as a vehicle for delayed release of protein drugs. The incorporation of biocompatible PEG can protect the protein from degradation, and subsequently, their bioavailability. PEGs are used for improving the blood compatibility of polymers and are capable of preserving the biological properties of proteins.37,45 Thus it seems that the chitosan/PEG–alginate system is a good candidate for oral delivery of proteins and other biomolecules.

11.3

Biomaterials for tissue engineering

In recent years, much attention has been focused on the development of biomaterials for use in tissue engineering applications. The increasing shortage of tissue and organ donors has driven the development of viable tissue substitutes. Tissue engineering is a multidisciplinary field of biomedical research that merges the fields of cellular and molecular biology, cell and tissue culture techniques, and materials science to recreate tissue for treatments involving reconstruction or replacement. New biocompatible materials are developed constantly for use as scaffolds in the tissue engineering of a wide variety of tissues and structures, such as cartilage, bone, liver, blood vessels, heart valves, myocardial tissues and the knee meniscus.47–49 Tissue engineering is defined as the creation of new tissue for the therapeutic reconstruction of the human body, by the deliberate and controlled stimulation of selected target cells through a systematic combination of molecular and mechanical signals.47,48

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In general, depending on the tissue that is to be engineered, two main tissue-engineering approaches can be distinguished. In the first and currently the most popular approach, cells are seeded onto scaffolds and cultured in vitro to form a construct that subsequently is implanted into a laboratory animal.50 The concept is to take a suitable material that can give physical form or outline to the area of tissue to be regenerated and to invest it with those tissue components that are responsible for tissue growth and repair. The combination, often called a construct, is able to help the patient regenerate new functional tissue. The tissue components may be cells or biomolecules or both. On the basis of this analysis, a definition of tissue engineering has been produced as follows – tissue engineering is the persuasion of the body to heal itself through the delivery to the appropriate sites of molecular signals, cells, and /or supporting structures.51 The most common type of tissue engineering product is one that involves a biodegradable polymeric support (for example, in the form of a porous scaffold) into which is introduced cells of the appropriate phenotype, along with some suitable drugs such as growth factors. This construct may be cultured in a sterile bioreactor, when the cells produce the regenerated tissue, which can then be implanted in the host.50–52 This is a very ambitious objective and only limited success has been achieved so far in relatively simple areas, such as the skin. Developments are well advanced in bone and cartilage regeneration, and there is much activity in experimental systems for arteries and nerves. In the second approach, which is gaining in popularity, especially with the advent of so-called smart scaffolds,50 acellular scaffolds are implanted into a laboratory animal. Rather than relying on in vitro seeding of cells as in the first approach, this alternative approach relies on in vivo ingrowth of surrounding tissue into the scaffold, with subsequent proliferation and differentiation into the desired tissue. Especially in large three-dimensional (3D) scaffolds, with a size exceeding the critical diffusion distance, this approach will be useful when ingrowth of vascularized tissue can be induced.52–55 The development of new biomaterials for tissue engineering, and the enhancement of tissue ingrowth into existing materials by means of attaching growth factors (‘smart scaffolds’), creates the necessity to develop tools for assessment of tissue ingrowth rates into porous biomaterials. A variety of materials has been considered for the scaffolds,53–58 the most common being the biodegradable polyesters mentioned previously, and some natural biopolymers such as collagen and a variety of polysaccharides (Table 11.9). Biodegradable synthetic polymers offer a number of advantages over other materials for developing scaffolds in tissue engineering, drug delivery and in vivo sensing.50 Poly lactic acid, poly glycolic acid, and their copolymers, dominate as scaffold materials because the materials are

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Table 11.9 Scaffold materials used for tissue engineering applications Synthetic biodegradable polymers (PLA, PLGA, PGA, etc.) Synthetic non-biodegradable polymers (as hybrids) Natural biopolymers (proteins, polysaccharides, silk) Self-assembled biological structures Tissue-derived structures (small intestinal submucosa, porcine pericardium, porcine dermis, amniotic membrane, etc.) Bioactive ceramics and glass ceramics Composites, including nanocomposites Multilayered structures Temperature-/pH-responsive polymers (cell sheet)

already approved by the FDA and because the degradation rate can be manipulated. However, the materials are much weaker and much less elastic than soft biological tissue, cell adhesion is only achievable with surface modifications, and the degradation rate is very fast. Other classes of degradable polymers have been investigated, such as polycaprolactone, polyanhydrides, polyphosphazenes, and so on.55–59 Recent developments have proposed that copolymers of polyethylene glycol with PLA or PLGA and the elastomeric polyglycerol sebacate should be examined for soft tissue applications.60

11.3.1 Surface engineered biomaterials for tissue engineering Several extracellular matrix (ECM)-like materials that combine synthetic polymer with three-dimensional collagen gels have been investigated as tissue engineering scaffolds. For example, polyethylene glycol–collagen composites have been used as in vivo scaffolds for connective tissue regeneration,61 whereas fibronectin–collagen and laminin–collagen composites have been used to grow cells in vitro. Functional receptor-mediated and signal-transmitting cell adhesion on a conventional biomaterial is mediated by ECM molecules, such as fibronectin, vitronectin, collagen or laminin. Numerous studies have highlighted the importance of ECM components for appropriate cell migration during morphogenesis of the nervous system, as well as for neurite outgrowth in vitro.61,62 The ECM proteins such as laminin (LN) and fibronectin (FN) are expressed in distinct spacial and temporal patterns in the developing nervous system. These proteins are necessary for natural healing processes of the nervous system, and appear to be a critical factor in the integration

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and function of biomaterial-based implants in the nervous system. Chandy and Rao63 used a spray coating technique to prepare poly(lactic acid) (PLA) tubes. To improve the flexibility of these devices, an elastomeric polymer, poly(ethylene vinyl acetate) (PEVAc), was added to the PLA. The PLA/ PEVAc tubes were further surface modified with elastin and laminin, via carbodiimide and glutaraldehyde treatment. This study63 indicated that the surface grafted matrix components (elastin and laminin) on PLA/ PEVAc tubes would offer a new approach for nerve growth and tissue engineering. The two main classes of extracellular macromolecules that form the natural ECM are proteoglycans and fibrous proteins. Tan et al.64 suggested that proteoglycans, containing long unbranched polysaccharide side chains covalently tethered to a fibrous backbone, form three-dimensional networks of hydrated gels in which cells can be embedded. The development of suitable three-dimensional matrices for the maintenance of cellular viability and differentiation is critical for applications in tissue engineering and cell biology. To this end, gel matrices with different proportions of alginate/elastin/polyethylene glycol (Alg/Ela/PEG) were prepared and examined.65 The composite matrix of Alg/Ela/PEG has polysaccharide structures (Alg) and fibrous proteins (Ela) in a water soluble PEG. This structure allows the Alg/Ela/PEG gel to maintain structural integrity and can keep biological integrity of cells to grow and multiply in the scaffold. This novel composite matrix structure resembles natural ECM components and may have potential biological and mechanical benefits for use as a cellular scaffold. Tissue engineering generally involves combining mammalian cells (including stem cells) with polymer materials, in such a way as to create new tissues or organs. Although this area has been the focus for considerable research,52,53 many challenges remain. First, it will be essential to find cell sources that yield sufficient quantities of differentiated cells. Stem cells may represent an important source, provided that their differentiation, growth and immunogenicity can be controlled. Another possibility is autologous cells, although in many cases one may not be able to grow them quickly enough and in such a way that they maintain their differentiated state. A second challenge is ensuring that cells in engineered tissues survive. One promising approach is to develop methods for vascularizing threedimensional cell scaffolds via microfabricated systems or the controlled release of growth factors.55,56 A third major issue is immune rejection. Finding ways to prevent it – for example, through the use of somatic nuclear transfer – is clearly important but far from resolved in practice. Finally, there are practical issues, such as the shelf-life of polymer–cell systems, as well as scale-up and production issues. Despite these many challenges, the opportunities in tissue engineering are numerous.

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Role of the scaffold and loaded drug/growth factor in the integration of extracellular matrix and cells at the interface

The molecular mechanisms that determine survival, differentiation and movement in multicellular organisms are dependent on interactions with the ECM. Cells in tissues are structurally and functionally integrated with their surrounding ECM via numerous dynamic connections. On the intracellular face of these linkages, adhesion receptors tether the contractile cytoskeleton to the plasma membrane and compartmentalize cytoplasmic signaling events, while at the extracellular face, the same receptors direct the deposition of the ECM itself.61,66 Such membrane–proximal functions, in turn, trigger distal processes, such as alterations in the direction of cell movement and the regulation of cell fate, and the construction of ECM networks and consequent shaping of higher-order tissue structure. An understanding of the molecular events that underpin ECM function would therefore help elucidate some of the key organizing principles of multicellular life. ECM and cell–ECM interactions also contribute widely to disease. Many of the major human diseases are either caused by defects in cell–ECM coordination, are exacerbated by aberrant use of normal cell adhesive processes, or are potentially correctable by altering tissue structure or cell movement.61,62,66 For example, progressive extracellular remodeling in chronic atherosclerotic, fibrotic and neurodegenerative diseases, leads to a loss of tissue integrity, altered adhesion is a defining characteristic of malignancy, and the pathogenesis of inflammatory and thrombotic diseases relies on aberrant cell aggregation and/or trafficking. The development of strategies to correct ECM dysfunction therefore has enormous promise as a route for improving treatment of many important clinical conditions. Growth factors initiate and control a variety of cellular processes involved in tissue formation (Table 11.10). Their use in the clinic, however, has been facilitated following advances in recombinant protein technology. Growth factors, growth factor receptors and monoclonal antibodies are currently being employed to treat clinical conditions such as obesity, cancer, and idiopathic short stature,55,56,66 with the potential for use in wound healing and tissue regeneration. Localized delivery of tissue inductive factors from scaffolds can function to direct progenitor cell differentiation toward the desired cell fate. Although in vitro studies of tissue formation on scaffolds can be performed simply by adding growth factors to cell culture media, translation of these studies to applications with in vivo tissue formation requires the use of delivery systems that can make these factors available at the appropriate concentration and duration.

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Biointegration of medical implant materials Table 11.10 Growth factors delivered to promote tissue formation Growth factor

MW/kDa

Functions

Nerve growth factor (NGF)

27

Insulin-like growth factor-1 (IGF-1)

27.9

Insulin-like growth factor-2 (IGF-2) Epidermal growth factor (EGF) Fibroblast growth factor-1 (FGF-1) Fibroblast growth factor-2 (FGF-2)

35.1

Promotes neuron survival and extension in the central and peripheral nervous systems; modulates differentiation of various neuron types in vivo and in vitro; role in tissue repair and fibrosis Mediates actions of growth hormone; increases proteoglycans and Type II collagen synthesis Promotes myogenic differentiation of ES cells Wound healing

6.2 17.5 17.3

Platelet-derived growth factor (PDGF)

22–25

Bone morphogenic protein-2 (BMP-2) Transforming growth factor-β1 (TGF-β1)

44.7

Vascular endothelial growth factor (VEGF)

19–22

25

Wound healing, vascular repair; fibroblast mitogen Chondrogenesis; angiogenesis; neuronal and endothelial cell proliferation Maturation of blood vessels, recruitment of smooth muscle cells to developing vasculature; wound healing; neural regeneration Osteogenesis; angiogenesis Promotes chondrogenic differentiation; increases cartilage matrix synthesis and chondrocyte proliferation Angiogenesis; vasculogenesis; osteogenesis; neurotrophic factor for motor neurons; cartilage remodeling

A tissue-engineered implant is a biologic–biomaterial combination in which some component of tissue has been combined with a biomaterial to create a device for the restoration or modification of tissue or organ function. Specific drugs/growth factors, released from a delivery device or from co-transplanted cells, would aid in the induction of host paraenchymal cell infiltration and improve engraftment of co-delivered cells for more efficient tissue regeneration, or ameliorate disease states. Growth factors are polypeptides that transmit signals to modulate cellular activities. Growth factors can either stimulate or inhibit cellular proliferation,

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differentiation, migration, adhesion and gene expression.66–70 There are several characteristic properties of growth factors. The characteristic properties of growth factors provide a biological basis for their use in tissue engineered devices. The principles of polymeric device development for therapeutic growth factors or drug delivery must be explored for tissue engineering applications.

11.4.1 Induction of angiogenesis in tissue-engineered scaffolds for bone repair: a combined gene therapy–cell transplantation approach One of the fundamental principles underlying tissue-engineering approaches is that newly formed tissue must maintain sufficient vascularization to support its growth. Efforts to induce vascular growth into tissue-engineered scaffolds have recently been dedicated to developing novel strategies to deliver specific biological factors that direct the recruitment of endothelial cell (EC) progenitors, and their differentiation.68,70 The challenge, however, lies in orchestration of the cells, appropriate biological factors, and optimal factor doses. Beck et al.68 reported a novel approach as a step forward to resolve this dilemma by combining an ex-vivo gene transfer strategy and EC transplantation. The utility of this approach was evaluated using threedimensional (3-D) poly(lactide-co-glycolide) (PLGA) sintered microsphere scaffolds for bone tissue engineering applications. Adipose derived stromal cells (ADSCs) were isolated and transfected with adenovirus encoding the cDNA of vascular endothelial growth factor (VEGF). They demonstrated that the combination of VEGF-releasing ADSCs and ECs results in marked vascular growth within PLGA scaffolds and thereby delineate the potential of ADSCs to promote vascular growth into biomaterials.68 Similarly, delivery of multiple growth factors to sites of bone injury was shown to dramatically enhance bone regeneration. Dual delivery of bone morphogenic protein-2 (BMP-2) and transforming growth factor-β3 (TGFβ3) from a hydrogel promoted significant bone formation by cotransplanted bone marrow stem cells (BMSCs) within six weeks of implantation.66 Interestingly, the synergistic activity of these two factors allowed them to be provided at a low dose, whereas supraphysiological concentrations of the individual factors resulted in negligible bone tissue formation. In addition to dual protein delivery, DNA and protein delivery can be combined to provide the necessary factors for tissue formation. Combined delivery of plasmid DNA encoding for BMP-4, VEGF protein, and human bone marrow stromal cells (hBMSCs) significantly enhanced bone formation relative to delivery of any single factor.

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11.4.2 Hydrogel composite materials for enhanced neurotrophin delivery in neural prostheses Long-term implanted electrode arrays have a significant importance in neural prosthetics for recording the electrical signals from nearby neurons and generating electrical signals to stimulate nearby tissue. Therefore, the density of neurons and their proximity to the electrode sites play an important role in the electrode performance.69 Many biodegradable and biocompatible polymers have been used as electrode coatings to minimize the acute and chronic inflammatory response and to preserve neurons near the recording sites. Since poly(ethylene glycol) (PEG) is non-toxic, nonimmunogenic, and inert to most biological molecules such as proteins, block copolymers of PEG have been studied by various groups36–39,44,69 to form hydrogels as the scaffolds that can support or stimulate neuron growth. A composite drug delivery system comprising polymeric hydrogel (i.e. poly(ethylene glycol)-poly(lactic acid), (PEGPLA) or poly(ethylene glycol)-polycaprolactone, (PEGPCL)), and other vehicles, such as biodegradable poly(lactic co-glycol acid) (PLGA) microspheres or polycaprolactone (PCL)-based electrospun nanofibrous scaffolds, was developed to provide neurotrophins at a predetermined rate for 2 ∼ 3 months in vivo.69 The stable and sustained release of neurotrophins can remarkably enhance the attraction, attachment and restoration of neurons around the chronically implanted electrodes.

11.4.3 Fine-tuning Notch signaling to promote angiogenesis Tissue engineering scaffolds releasing proteins mimic the natural reservoir capacity of the ECM. Coupling delivery of protein with a degradable carrier capable of organizing tissue formation is a powerful approach for tissue regeneration. Promoting angiogenesis is both a critical aspect for tissue regeneration and a potential effective therapy for cardiovascular diseases. Inspired by the recent findings that systemic introduction of Notch inhibitors reduced blood flow to tumor tissues by forming excessive yet dysfunctional vasculature, various studies have been conducted70,71 and have proposed that precisely controlled partial and local Notch inhibition might enhance regional neovascularization, by altering the responsiveness of local endothelial cells and/or progenitors to angiogenic stimuli. In vivo, delivery of an appropriate combination of gamma secretase inhibitor (GSI) and VEGF from an injectable alginate hydrogel system to ischemic hind limbs led to a greater recovery of blood flow than VEGF or GSI alone; perfusion levels reached 80% of the normal level by week 4 using combined GSI and VEGF delivery.71,72

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Strikingly, direct intramuscular or intraperitoneal injection of GSI did not result in the same level of improvement, suggesting that the extended presence of GSI (gel delivery) is important for its activity. The optimal dose of GSI delivered from alginate hydrogels did not show any adverse effects, in contrast to systemic introduction of GSI. Altogether, these results suggested a new approach to promote angiogenesis by fine-tuning Notch signaling, and may provide new options to treat patients with diseases, such as diabetes, that can diminish angiogenic responsiveness.72

11.4.4 Myocardial tissue engineering via growth hormones The concept that growth hormone (GH) and insulin-like growth factor 1 (IGF-1), the mediator of many of the effects of GH on peripheral tissues, target the heart has recently emerged from a series of animal and human studies. Conditions of GH/IGF-1 deficiency in humans are associated with cardiac atrophy and impaired cardiac function. The hypertrophic response with enhanced cardiac performance observed in rats subjected to chronic GH/IGF-1 excess appears to be beneficial in the setting of experimental and, recently, human heart failure.73,74 It is likely that the site-specific delivery of GH or IGF-1 via polymeric scaffolds can regenerate the cardiac tissues through cell proliferation and differentiation.72,73 Substantial data suggest that IGF-1 is a potent cardiomyocyte growth and survival factor.73 Mice deficient in IGF-1 have increased apoptosis after myocardial infarction, whereas cardiac-specific IGF-1 over expression protects against myocyte apoptosis and ventricular dilation after infarction.66 IGF-1 over expression increases cardiac stem cell number and growth, leading to an increase in myocyte turnover and function in the aging heart. After infarction, IGF-1 promotes engraftment, differentiation, and functional improvement of embryonic stem cells transplanted into myocardium. Davis et al.73 developed a delivery system using a ‘biotin sandwich’ approach that allows coupling of a factor to peptide nanofibers without interfering with self-assembly. Biotinylation of self-assembling peptides allowed specific and highly controlled delivery of IGF-1 to local myocardial microenvironments, leading to improved results of cell therapy. This approach allowed a greater control of the intramyocardial environment by delivering growth factors to injured myocardium. With this system, it may be possible to design the local microenvironment to improve the endogenous regenerative response; for example, by delivery of a chemoattractant to promote stem cell migration. Although much more must be learned about why mammals have inadequate cardiac regeneration, the ability to control the local myocardial microenvironment may prove critical to preventing heart failure.

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11.4.5 Multiple factor delivery for vascular tissue engineering Tissue morphogenesis and regeneration are typically driven by the concomitant action of multiple factors, which can work synergistically on the same process or can target different barriers to regeneration. The synergistic effect of growth factors has been reported for many developmental processes, including angiogenesis, where mature blood vessels form by the combined action of VEGF and platelet-derived growth factor (PDGF) to form stable vessels.66,73 Although VEGF is able to initiate angiogenesis, PDGF promotes vessel maturation via recruitment of smooth muscle cells to the developing endothelium. Poly lactic acid/poly glycolic acid scaffolds releasing both VEGF and PDGF formed a mature vascular network within and around the scaffolds. The concentration and duration of function for tissue inductive factors at the regenerating tissue site are critical parameters involved in promoting developmental processes and the formation of mature tissues. Therapeutic angiogenesis and anti-angiogenesis reveals these concepts, as immature vessels or vessels that regress over time can lead to unsuccessful or abnormal tissue formation. Sustained expression of low to medium levels of VEGF or hormone combinations are required to promote the growth of blood vessels displaying normal morphological and functional characteristics.66–69 Transplanted cells that expressed low levels of VEGF avoided the formation of aberrant vessels and hemangiomas observed with cells expressing high levels of VEGF. Thus, it is assumed that multi-factorial presentation of growth factors can be more effective at stimulating natural developmental processes leading to tissue formation.

11.5

Future trends

Combination products that converge medical devices with therapeutics and diagnostics are a rapidly growing segment of the health care industry. Combination devices – those comprising drug releasing components together with functional prosthetic implants – represent a versatile, emerging clinical technology promising to provide functional improvements to implant devices in several classes.75,76 Landmark antimicrobial catheters and the drug-eluting stent have heralded the entrance, and significantly, routes to FDA approval, for these devices into clinical practice. Most prominent are new combination devices representing current orthopedic and cardiovascular implants with new added capabilities from on-board or directly associated drug delivery systems which are now under development. Wound coverings and implantable sensors will also benefit from this combination enhancement. Infection mitigation, a common problem with implantable

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devices, is a current primary focus. Ongoing progress in cell-based therapeutics, progenitor cell exploitation, growth factor delivery and advanced formulation strategies will provide a more general and versatile basis for advanced combination device strategies.76,77 These seek to improve tissue– device integration and functional tissue regeneration. While the field of biomedical engineering encompasses much more than combination products, many of today’s most promising areas of development combine biologic, pharmaceutical, and/or medical device components. Among the most notable biomedically engineered combination products are neuro-modulating devices, tissue-engineering technologies, and nanomedicines.76 The drug eluting stents significantly reduce coronary in-stent restenosis compared with bare metal stents.78 However, recent studies have shown that late stent thrombosis, defined as thrombosis after 30 days of post-deployment, is a growing safety concern with drug eluting stents.78 The drug eluting stents did not fully endothelialize after 40 months of implantation, whereas the bare stents can completely endothelialize after 6 to 7 months.76,78 This impaired initial healing of drug eluting stents may be caused by the polymer matrix, the drug, or a combination of the two.78 Recent studies have indicated that surface engineering of drug eluting stents with extra cellular components (such as collagen and laminin) can enhance the healing and endothelialization faster.78,79 Thus it seems that a combination of drug release with surface engineering stents may be a novel way of neointima formation and naturalization. Future combination devices might best be completely re-designed de novo to deliver multiple bioactive agents over several spatial and temporal scales to enhance prosthetic device function, instead of the current ‘add-on’ approach to existing implant device designs never originally intended to function in tandem with drug delivery systems.

11.6

References

1. anderson jm. Mechanisms of inflammation and infection with implanted devices. Cardiovasc. Pathol. 1993:2:33S–41S. 2. singer i, hutchins gm, mirowski m, mower mm, veltri ep, guarnieri t, griffith lsc, watkins l, juanteguy j, fisher s, reid pr, weisfeldt ml. Pathologic findings related to the lead system and repeated defibrillations in patients with the automatic implantable cardioverter defibrillator. J. Am. Coll. Cardiol. 1987:10:382–388. 3. joist jh, pennington dc. Platelet reactions with artificial surfaces. Trans. Am. Soc. Artif. Intern. Organs. 1987:4:33–341. 4. vroman l. The life of an artificial device in contact with blood: Initial events and their elect on its funal state. Bull. NY Acad. Med. 1988:7:64–352. 5. noishiki y, chvapil m. Healing pattern of collagen-impregnated and preclotted vascular grafts in dogs. Vasc. Surg. 1987:11:21–401.

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12 Mechanisms of failure of medical implants during long-term use A. K A S H I and S. S A H A, SUNY Downstate Medical Center, USA

Abstract: This chapter discusses the mechanisms that are responsible for implant failure following clinical use. The chapter focuses mainly on the mechanical and biological factors (for both hard and soft tissue implants) that can influence the in vivo survival of implants. Techniques including microscopy and finite element analysis, to study the stress distribution of implants; design considerations to improve future generations of implants are also discussed. Key words: implant failure, fatigue, hard and soft tissue replacements, wear, metallic implant, non-metallic implant.

12.1

Introduction

Medical implants are routinely used to rehabilitate patients with lost or damaged tissues. The main goal of implants is to help patients with disabilities to return to normal function for the longest possible duration. Further, implants can be used either to augment existing performance of the body or to replace missing tissues, organs or parts of the body. The interplay of several in vitro and in vivo factors is responsible for the clinical success of implants. However, similar to other artificial materials, implants have a finite lifespan, typically following long-term use. Load-bearing implants/ prostheses are typically subjected to fatigue following millions of cycles of use, which along with other causes can lead to their failure (Zhang et al., 2004; Flanagan et al., 2008; Kim et al., 2007; Chee and Jivraj, 2007; Sudhakar, 2005). These might also include aseptic loosening of prostheses (Unwin et al., 1996; Malchau et al., 1993). Similarly, non-load-bearing implants can fail after certain durations due to various etiologies including infections, immune suppression and/or physiologic loss of surrounding natural tissues. Although in vitro studies are useful to study medical implants, their extrapolation to in vivo performance remains largely unproven. A multidisciplinary approach involving biomedical engineers, clinicians and researchers is critical for the investigation of failed implants. In addition, animal experimentation needs to be encouraged to identify the biological response to new biomaterials that are being developed for the manufacture 326 © Woodhead Publishing Limited, 2010

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of implants. Adopting recent scientific techniques, including evidence-based approaches, is also useful to study implant survival rates and treatment outcomes. Studying failed implants will help researchers and clinicians to understand the causal factors for failure and to develop better iterations of prostheses for future use. Typically, this includes gross and microstructural analysis. Analytical techniques, including histopathology, scanning electron microscopy and finite element analysis, are useful for examination of failed implants and implant behavior. Further, this chapter will discuss mechanisms (e.g. mechanical and biological) that can contribute to implant failure following long-term use.

12.1.1 Failure mechanisms of medical implants Medical implants can fail by various mechanisms, which can include either single or multiple etiologies. From a biomechanical perspective, it is useful to note that there generally exist two different types of implants – loadbearing and non-load-bearing. The failure mechanisms associated with these two types can vary, depending on their anatomical location, the loads that they are subjected to during function, and other associated factors including age, sex, and the general systemic health of a patient. Typically, in the case of load-bearing implants, the generation of wear debris is a critical factor that can contribute to foreign body reactions and metallosis. This is seen mostly in the case of orthopedic hip and knee implants. Ceramic implants are known to cause lesser wear debris when compared to their metallic counterparts (Rahaman et al., 2007). This is due to better tribological properties of bioceramics (the possibility to produce a smoother finished surface, thus leading to lesser friction during use), and superior biocompatibility when compared with metals. Apart from metallic implants, authors have studied debris resulting from the wear of failed elbow, wrist and finger silicone implants (Hirakawa et al., 1996). In the case of non-load-bearing implants (e.g. maxillofacial implants, including chin, cheek and cranial prostheses, eye and ear implants and breast implants), the failure mechanisms might be different. For instance, authors have previously reported breast implant failures from folds/flaws, although this is not the most dominant form of failure for these types of prostheses (Brandon et al., 2006).

12.2

Manufacturing deficiencies

The manufacture of an implant typically requires strict adherence to guidelines that have been developed by medical device makers following good manufacturing practices (required by the Food and Drug Administration – FDA). Failure to follow these prescribed guidelines can compromise the structural stability of implants. Further, several cases of implant failure have

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been shown to result from deficiencies in the finished product due to manufacturing defects. These can result as a consequence of improper material selection, poor quality control, or a lack in oversight during handling (e.g. sterilization and packaging). Authors have previously reported implant failures that can be attributed to impurities in the raw materials. These impurities can act as crack initiation sites to allow for further deterioration of the entire implant structure during long-term use. Failures of the past, most notably the Vitek temporomandibular joint (TMJ) implant and the Dow Corning Silicone breast implant failures, have brought to the forefront issues related to appropriate material selection and adequate pre-clinical testing (Ta et al., 2002; Westermark et al., 2006; Mercuri and GiobbieHurder, 2004; Mercuri and Anspach, 2003; Speculand et al., 2000; Kearns et al., 1995). These includes in vitro laboratory testing of implants and animal studies. Additionally, the need for liability protection for biomaterial manufacturers and tort reform are important issues that need further discussion. Biomedical device manufacturers are required to follow best manufacturing practices in order to ensure that the finished products (i.e. implants) conform to industry, physician and patient requirements. Since an implant might be subjected to contamination and/or deterioration during its manufacture, it is important to characterize the possible failure mechanisms during the manufacturing process.

12.3

Mechanical factors (e.g. fatigue, overloading)

Factors that might influence failure of an implant include the choice of biomaterial, the overall geometry of the prostheses, the magnitude of forces and the number of cycles of use that it is subjected to. Typically, implants are designed to withstand millions of cycles of use within the in vivo environment. However, in vitro testing and evaluation might sometimes not be sufficient to predict the lifetime survival rate due to secondary contributing factors including immune suppression, infections, and other secondary illnesses. Almost all load-bearing implants (e.g. hips, knees, dental) are subjected to a combination of axial (i.e. compressive and tensile), bending and torsional forces during function. It is important to note that high stresses would most likely result in single-cycle and low-cycle implant failures, while low stresses (i.e. high/multiple cycles that typically last a million or more in number annually) would mostly lead to fatigue types of failure. This behavior is especially important in the context of metallic implants (Geesink et al., 1988). From a biomechanical perspective, the likelihood of failures can be minimized if the extent of the applied loads (stresses) can be reduced. It is important to note that most orthopedic implants fail from fatigue if bone does not heal. Some of the reasons for non-healing in skeletal tissues

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include excessive motion, lack of blood supply, infection, immune suppression, systemic illnesses, and biomechanical factors such as stress shielding of implants. If an implant is subjected to stresses that are beyond its endurance limit, the structure will ultimately fracture/fail. However, if an implant is subjected to low stress levels, it can withstand millions of cycles of use, although these repeated insults can create microscopic fracture zones that can contribute to larger structural failures in the long-term, due to fatigue. This type of fatigue failure results because of localized stress concentrations on certain points of an implant before other areas of the structure are affected. This can initiate a crack that can eventually propagate to involve the entire implant structure. Compressive, tensile and torsional loads are simultaneously acting on a load-bearing implant during function. In order to fully elucidate the failure criteria of implants, it is essential that biomechanical evaluation incorporates tests that can determine its mechanical properties (i.e. compression, tension and torsion). We have previously reported the pull-out resistance of dental implants. The objective of the study was to compare the pull-out resistance of small and large diameter (3.25 and 4.5 mm) dental implants, and the relationship of these implants to bone density. Two groups of implants, consisting of eighteen implants of each diameter, were placed in five embalmed human mandibles. The bone mineral density of the area surrounding the implant site in the coronal cross-section was measured by quantitative computed tomography (QCT). The initial implant stability was tested with a periodontium diagnostic device and the pull-out resistance was tested with a mechanical testing system. Results showed the same initial stability for the two implants. However, the maximum pull-out force required for the large diameter implants was 15% greater than that required for the small diameter implants (Kido et al., 1997). Further studies need to be carried out to statistically validate these results. The fatigue behavior of biomaterials is generally characterized by measuring the applied stress versus number of cycles (S–N curve) for failure (Fig. 12.1). Although in vitro testing and evaluation is helpful in obtaining values for strength, hardness, fracture toughness and wear, these tests might not be sufficient to predict their lifetime survival rates due to secondary contributing factors, including age, sex, body mass index (in the case of orthopedic hip and knee implants), immune suppression, infections, and other secondary illnesses.

12.4

Wear

Wear is a common problem with artificial implants, especially for total or partial joint replacements. In the case of metallic implants, particulate disease resulting from wear debris is a frequently encountered clinical

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Stress, s

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Number of cycles, N

12.1 S–N curve showing increase in fatigue life as a consequence of decreased stresses.

problem that can sometimes warrant revision surgery. The problem of wear is often seen in total hip and knee replacements. We have previously reported findings of catastrophic peri-implant bone loss caused by polyethylene and metallic wear in total knees in two cases. The first was a pathologic fracture of the distal femur that was associated with catastrophic polyethylene failure, while the second was related to three-body wear of the polyethylene from detached CoCr beads (Gustafson et al., 1993). Many authors have studied wear behavior of total joint replacements and other implants previously (Wright et al., 1992; Sumner et al., 1995; Revell, 2008; Revell et al., 1997; Fehring et al., 2004; Marshall et al., 2007; Khan et al., 1996; Hirakawa et al., 1996; Korovessis et al., 2006; Agins et al., 1988; Yamaguchi et al., 1997; Griffin et al., 2007; Ingham and Fisher, 2000; Howie et al., 1988; Sieber et al., 1999; Schmalzried et al., 1996; Al Jabbari et al., 2008). Figure 12.2 is a retrieved polyethylene liner (from a knee implant) showing wear facets. Some of the observations from gross examination of this retrieved prosthesis include delamination, excessive wear and partial fracture at the edge. Although the issue of wear is of major concern with load-bearing implants, when compared to metals, ceramic implants are known to have better tribological properties due to their ability to be polished to very smooth surfaces, thus reducing friction and consequently producing less wear during function. However, ceramics are prone to other problems during use. For instance, the formation of cracks following repeated use can act as initiation sites that can progress to complete failure of a ceramic implant. The phenomenon of crack initiation and propagation has been studied extensively with respect to ceramic dental crowns (Zhang and Lawn, 2005; Bhowmick et al., 2005; Zhang et al., 2004, 2008; Kim et al., 2007). This knowledge might

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5.00 mm

12.2 Polyethylene liner from a knee implant retrieved from patient, showing wear facets.

also be useful in understanding flaws and their effects on the structural stability of other orthopedic implants, including artificial hips and knees. Once there is a critical flaw in a brittle material such as a ceramic, the energy needed to drive it further into the structure of the material becomes much lower. Biological fluids in vivo can accelerate crack growth in ceramics by reducing the energy needed for crack propagation. When this is coupled with a lack of conduits (in an implant structure) for the escape of biological fluids, it can increase the hydrostatic pressure within this crack region. If these previously mentioned phenomena are combined with cyclic loading, which is commonly seen in load-bearing implants, it might accelerate crack growth, leading to catastrophic failure. Although ceramics are known to withstand compressive loads, they can easily fail when subjected to tensile loads. The lack of tensile strength makes them less attractive for many loadbearing applications. Another emerging complication with ceramic hip prostheses is the generation of acoustic emissions from stripe wear. Although this has been only anecdotally reported, by a few investigators, one of the suggested theories for this occurrence has been attributed to edge loading between the acetabular component and the corresponding ceramic ball (Fig. 12.3). Cases of similar nature have been reported in the medical literature and have resulted in their removal from patients (Taylor, 2006; Restrepo et al., 2008; Walter, 2007). Figure 12.4 shows an alumina hip implant that was in situ for 2 years in a 54-year-old patient before it had to be removed due to a persistent squeaking sound during function. On examination, it was observed that the ceramic ball and its corresponding acetabular component showed

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15:02:15

12.3 Stripe wear on retrieved ceramic hip implant (alumina femoral head).

12.4 Retrieved alumina hip implant (acetabular component); black arrow shows probable location of edge loading.

wear facets. This is seen in squeaking ceramic hip prostheses. The alumina femoral head (with stripe wear as a result of titanium wear debris in situ) and acetabular component are shown in Figs 12.3 and 12.4, respectively. It is also evident from the acetabular component (area where arrow is shown in Fig. 12.4) that there appears to be some surface damage to an area

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corresponding to the edge. This might support the theory that edge loading can be a contributing factor to the audible squeak. Since ceramic hip implants have been used for only a relatively short time in the US, when compared to the more traditional metal-on-metal or metal-on-polyethylene types, clinical complications that have occurred with ceramic implants have been reported only in the past two to three decades. It will be important to study the failure mechanisms of these newer implant types as they might be used in greater numbers clinically in the future, especially in younger, active adults. This may be attributed to improved material properties and designs, as well as better long-term clinical results. Since younger patients are more active when compared to the elderly, studying the correlation between age and level of physical activity that can produce specific types of implant wear patterns might be useful. Moveover, changes in the surface properties and microstructural features of the ceramic material need to be studied in order to further delineate the underlying causes for these phenomena.

12.5

Corrosion

Corrosion is a complication that can have multiple detrimental effects, both clinically and from a materials standpoint. This is especially important in the context of metallic implants. Examples of some of the types of corrosion include pitting, erosion, galvanic/two-metal corrosion, leaching, stress corrosion and crevice corrosion. Authors have reported that although the rates of corrosion are less prevalent than in the past, they continue to pose a clinical problem with currently used metal implants (Hallab et al., 2001, 2004, 2006; Upadhyay et al., 2006; Rose et al., 1972; Brown et al., 1995; Sutow et al., 1985; Khan et al., 1996; Urban et al., 1994; Harding et al., 2002; Reclaru and Meyer, 1994; Case et al., 1994; Korovessis et al., 2006; Okazaki et al., 1998). In the case of metallic implants, especially modular types that have one or more components with tight tolerances, there can be higher chances for crevice corrosion to occur. Crevice corrosion can also be witnessed at regions between an implant and adjacent bone. Hard tissue implants are meant to be rigid when they are placed in vivo. However, it is well known that there is micromotion during function. These micromotions can lead to the generation of microscopic particulate debris and other toxic products, thus lowering the pH of the local environment. Implant motion that is greater than 150 μm leads to fixation by formation of mature connective tissue (Pilliar et al., 1986). Animal and human studies have confirmed these findings (Stulberg et al., 1991; Sumner et al., 1995). Studies have shown that the local pH in corrosion zones of an implant can reach values as low as 2.5 (Hallam et al., 2004). Due to high tolerances between mating compo-

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nents in an implant, biological fluids might not be able to fully enter the crevice regions to flush out these toxic products, leading to corrosion in the long-term. X-ray spectrographic microanalysis of failed implants due to corrosion can reveal the elemental composition of these corrosion byproducts in an implant.

12.6

Clinical factors for implant success and failure

12.6.1 Health of patient Implant success or failure depends to a great extent on the systemic health of the individual. For instance, it is well documented that patients with suppressed immune systems (e.g. long-term steroid users) and diabetics are prone to higher rates of infections when compared with their normal counterparts (McCracken et al., 2000; Moy et al., 2003; Takeshita et al., 1997, 1998; Beikler and Flemmig, 2003; Mombelli and Cionca, 2006; Gerritsen et al., 2000). Studies have shown that there is less osseointegration with titanium implants in diabetic animals (McCracken et al., 2000). However, these results are conflicting as there are other studies that have reported good fixation of percutaneous implants in diabetic animals (Gerritsen et al., 2000). A suppressed immune system is typically more prone to attack by opportunistic infections (e.g. candidiasis) and the presence of a foreign object such as an implant in such cases might lead to higher chances of failure. Since the average life-span of individuals is increasing in the US and many other countries, due to access to improved medical care, the percentage of surviving diabetics in geriatric sub-groups of the population has also been increasing. According to the Centers for Disease Control (CDC) in US, in the year 2007 the percent of non-institutionalized persons with diagnosed diabetes in the age group of 65–74 years was 19%, and individuals over the age group of 75 years comprised 18% (Prevention, 2007). Although it is difficult to predict the exact numbers, it is likely that a significant number of individuals in this age group will require implants. Further, the failure mechanisms of implants in the geriatric population might most likely result from higher chances of infection and other co-morbid factors due to poorer general systemic health. This highlights the need for further investigations, including long-term clinical observations and follow-up to study the survival rates of implants in these cases (Morris et al., 2000).

12.6.2 Surgical errors Surgeon proficiency is an important factor in implant surgery and its success. The placement of even a well-designed implant by an unskilled surgeon will

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have a high chance of failure. Factors including the surgeons’ level of experience with the placement of a particular type of implant, surgical technique and ease of handling of the armamentarium, are critical factors for implant survival outcomes (Shalabi et al., 2006). Failure to follow preoperative, intra-operative and post-operative protocols can contribute to implant failure. Esposito et al. mention that a combination of surgical trauma and anatomical conditions is probably the most important etiological factor responsible for early oral implant losses. In their review, they relate this to be approximately 3.6% out of 16 935 implants (Esposito et al., 1998). Simple steps, including changes in the orientation of an implant when it is surgically placed, can alter its wear properties. For instance, studies have shown that alignment of acetabular components of total hip prostheses that are greater than 50 degrees can lead to excessive wear of the polyethylene liners (Wright and Goodman, 2001). Similarly, changes in the orientation of other load bearing prostheses, including dental implants, can lead to excessive functional loads and accelerated localized bone resorption around them (Isidor, 1996, 1997). Implant design and surface finish are critical factors in their long-term clinical success and must take into consideration anatomy and biomaterials (Triplett et al., 2003; Sykaras et al., 2000; Misch, 1999; Beksac et al., 2006; Anitua, 2006; Esposito et al., 1998). Variations in implant designs, materials and shapes might produce different results, thus making it difficult to assume that similar wear patterns will be seen with all implants (Sykaras et al., 2000; Cehreli et al., 2004; Esposito et al., 1998). Therefore, further biomechanical studies are needed to better understand the different types of implant failures that might most likely be seen.

12.7

Failure mechanisms of non-load-bearing implants

Non-load-bearing implants (e.g. craniofacial prostheses) are generally not subjected to the same mechanical loads as their load-bearing counterparts (e.g. hip, knee and dental prostheses). Thus, the failure modes that might be seen with these implants are generally different. The issue of wear debris and particulate disease, which is particularly important in the case of loadbearing implants might not pose a significant threat toward failure in the case of non-load-bearing implants. However, other clinical complications, including secondary infections and bone resorption, are important factors that can lead to failures of non-load-bearing implants. Poly(methyl methacrylate) (PMMA) is used for cranial defect reconstruction due to its rigidity (unlike ceramics, including hydroxyapatite, which are brittle in nature) (Nassiri et al., 2009; Eppley, 2005; Eppley et al., 2004). However, there still exists scope for developing improved versions of implants; for example, biomaterials that can adapt to physiological

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growth (e.g. materials that possess the ability to conform to physiological craniofacial skeletal growth), especially in children. For instance, the development of materials that can promote osteoinduction, as well as being capable of resorption so that the defect(s) will be filled in with natural tissue at a desired time, will be useful. Such developments can be particularly helpful in craniofacial reconstruction in the pediatric population, where constant growth of the skeletal tissues is seen. In a biomechanical study by Eppley, different compositions of PMMA cranioplasty materials were evaluated by an impact resistance test (according to ASTM D 3029-78). The results of this study showed that the fracture patterns were different for porous versus solid implants. In the case of porous implants, a radiating stellate pattern originated from the point of impact. However, solid implants showed fracture patterns that were linear and nonstellate (Eppley, 2005). This experimental information might be useful in the design of future versions of biomaterials that take into consideration biomechanical factors including the force needed to fracture the prostheses and the likelihood of fracture in vulnerable populations, including children.

12.7.1 Soft tissue implants Soft tissue implants are generally used in the body for replacement of a variety of tissues, including ligaments, cartilages and cardiac valves, and for maxillofacial reconstruction. A few examples of soft tissue implants include silicone gel breast implants, polymeric chin, cheek and lip maxillofacial implants, bovine, porcine and human cadaveric cartilage, ligament and tendon replacements and bovine or porcine cardiovascular grafts for valve replacements. The biomechanical requirements of soft tissue implants are typically different when compared with implants used for hard tissue replacements. For instance, hip and knee implants are rigid structures, while most soft tissue replacements are viscoelastic in nature. Thus, their failure mechanisms (e.g. mechanical) will also be different when compared with hard tissue implants, although some common factors, including infections, can still lead to failures in any type of implant (Vinh and Embil, 2005; Ehrlich et al., 2005; Bauer et al., 2006; Neut et al., 2003; Widmer, 2001). Additional factors that might influence the survival outcomes of soft tissue implants include the general systemic health and co-morbid conditions (e.g. diabetes). Allogeneic as well as xenogeneic grafts have been known to cause foreign body reactions and rejections in some individuals. Similarly, silicone gel breast implants and polymeric maxillofacial implants have been associated with infections in some cases. Some rare failure mechanisms, for instance, the occurrence of fold flaw failures in breast implants have been previously reported by clinicians (Brandon et al., 2006). Other authors have studied

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platinum concentration in body fluids of women who have been exposed to silicone and saline breast implants (Lykissa and Maharaj, 2006). Authors have reported on the use of polyethylene, polypropylene, nylons, carbon fibers, polytetrafluoroethylene (PTFE) and polyesters as artificial ligament replacement materials. Most of these materials are known to fail clinically, despite their use in other anatomical locations more successfully (Amis, 2005). This might be related to the complex biomechanics of the knee joint and the effects of wear debris leading to adverse tissue reactions.

12.8

Failure analysis of medical implants

Implant failures can be studied by a variety of analytical/experimental techniques. Some examples include visual inspection, light microscopy, scanning electron microscopy (SEM), fractography and microhardness measurements. As a first step to study a failed implant, visual inspection can provide clues related to macroscopic fracture zones (Fig. 12.2). Although it might be difficult to quantitatively assess the failure criteria or to predict what might have occurred on a microscopic scale, visual inspection allows researchers sometimes to gather useful information, including fracture initiation and the directions of crack propagation. The next level of evaluation might include the use of light microscopy. This can highlight the loss of surface coatings on implants and the presence of any pathologic tissue. Figures 12.5a and 12.5b show close-up images of two retrieved knee implants that have been studied using a handheld digital microscope, while Fig. 12.5c shows the surface of a retrieved femoral head from a hip implant. For a higher resolution and to obtain a detailed microstructural analysis (e.g. grain size and distribution) of fracture zones and normal areas of failed implants, an SEM study will be useful. X-ray spectrographic microanalysis can provide information related to the chemical/elemental composition of retrieved implants. This will allow researchers to study changes in the chemical composition of failed implants, particularly being able to focus on failure and corrosion zones in a structure. In addition, this technique can be useful to analyze oxide layer compositions on implant surfaces, particularly metallic structures. The surface roughness of an implant is an important factor that can influence wear rate. Surface profilometry can be used to study the surface roughness of an implant. This characterization technique can allow researchers to analyze failed or retrieved implants. We have previously reported on the use of a non-contact surface profilometer to study and compare the surface roughness of a new and a retrieved TMJ implant (Fig. 12.6). The results of our study showed that the roughness parameters of retrieved implants were statistically higher (p < 0.05) than those for new specimens. Furthermore,

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(b)

(c)

0.500 mm

12.5 (a) Beaded surface of knee implant; (b) mesh-like appearance of retrieved knee implant; (c) surface of retrieved hip prostheses (metallic femoral head) showing scratch marks.

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12.6 Retrieved TMJ implant made of Co-Cr-Mo (Vitallium) alloy, showing large pits with irregular edges.

these roughness values were lower than those reported for retrieved metal total knee implants in the literature (Vaderhobli and Saha, 2007). Some of the other useful techniques for characterizing retrieved implants include microhardness and fractography. Newer characterization techniques that are capable of measuring hardness in the nanoscale range, by employing techniques including nanohardness indenters, and of studying surface topographies of implant structures using atomic force and threedimensional focus variation microscopes, are being used by researchers. They allow the study of materials and implant surfaces at much more detailed resolutions than would have been possible previously. In addition, characterization of implant surfaces at the regions of failure with these newer techniques can provide information related to local events occurring at the failure site. Software programs, including finite element analysis (FEA), can be useful to study or predict the behavior and probable failure mechanisms of implants/prostheses (Lin et al., 2007; Zhang et al., 2008). Useful information, such as response to fatigue, stress distribution in implant structures and design flaws, can be investigated using software simulation packages such as FEA. We have previously examined the stress distribution around the screws of a TMJ implant with applied loads using a FEA simulation package (Kashi et al., 2009; Roychowdhury et al., 2000). The results of our study showed that maximum stresses (von Mises) and strains were seen in the screw hole areas that were closest to the condyle region of the TMJ implant. Other simpler techniques that can be utilized to study contact stresses in joints include the use of pressure-sensitive film (Buechel et al., 1991).

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12.9

Conclusion

Authors have previously reported that biomaterial associated infections in orthopedics range from 1% to 3% (Van De Belt et al., 2001; Garvin and Hanssen, 1995). Peri-prosthetic infections have been reported to be as high as approximately 10% in the past in the field of orthopedics. However, with the use of modern surgical techniques (e.g. use of laminar air-flow operating rooms), these numbers have been minimized and are currently reported to be as low as 1% (Bauer et al., 2006). Although the majority of elective orthopedic procedures include replacement of hips and knees, it must be noted that other structures, including shoulder, ankle and craniofacial reconstruction, also fall under the category of orthopedic procedures. Artificial ligament and tendon replacements of the knee are also fraught with poor survival outcomes due to the complex biomechanical behavior of the knee. Tissue engineering efforts that might be able to focus on developing newer tendon and ligament replacements will be essential in the future. Developmental disorders (e.g. cleidocranial dysplasia and other congenital head and neck deformities), trauma and arthritic conditions affecting the maxillofacial region, including the temporomandibular joint (TMJ), warrant the need for developing better versions of implants to rehabilitate these structures. Since the use of TMJ implants is not as common as other implants (such as the hip and knee), their failure analysis has not been studied extensively. Important observations from other total joint replacements, including artificial hips and knees, can help researchers to better understand the behavior of newly developed biomedical implants. Efforts have been made to improve the survival outcomes of implants by minimizing post-operative infections. For instance, antibiotic-loaded bone cement is used by some surgeons during orthopedic implant placement to minimize infections, although there is no consensus among clinicians to standardize its use (Van De Belt et al., 2001). Further, there is a dearth of evidence-based studies to show if the effect of low dose antibiotics (from bone cements) leaching into the systemic milieu is detrimental, for instance, by promoting long-term antibiotic resistance. Studies have shown that the survival rate of dental implants in Type 2 diabetic patients is improved if chlorhexidine mouthrinse, an antimicrobial, is used at their time of placement (Morris et al., 2000). It might be interesting to study whether similar improvements in the survival rates of other orthopedic implants (e.g. artificial hips and knees) among Type 2 diabetics can be achieved with the use of chlorhexidine during their placement. Studies have shown that implant motion greater than 150 μm leads to fixation by formation of mature connective tissue (Pilliar et al., 1986). Animal and human studies have confirmed these findings (Sumner et al., 1995; Stulberg et al., 1991). Investigating this might help researchers to

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develop future iterations of prostheses with improved performance. As a consequence, studying the aforementioned issues will help in terms of providing a higher level of success for our future patients. The medical device industry has sales of at least US$ 100 billion worldwide and is highly research intensive (Ratner and Hench, 1999). For example, over a period of 12 years (from 1990 to 2002), the percentage of revenues reinvested in research and development in the medical device industry almost doubled. Consequently, along with this, investments during the past decades have led to encouraging results in terms of reduced incidence of mortality due to heart attacks, strokes, diabetes and breast cancer (Panescu, 2006). Implant survival can be improved from the combined efforts of clinical and laboratory (e.g. materials development) research. For instance, scientists are working to develop novel ceramic composite materials (e.g. zirconia-toughened alumina, ZTA) for artificial hips and knees, in order to counter phase transformation and aging phenomena in zirconia bioceramics (Chevalier, 2006). Researchers have attempted to characterize the microstructure of teeth in order to elucidate their remarkable resistance to fatigue and their ability to self-heal from microscopic fractures (Chai et al., 2009). Understanding this behavior might potentially be able to help scientists develop new-generation biomaterials with these characteristics. It is projected that the number of people in the US who are Medicare beneficiaries will increase to 75 million by the year 2030. This number was approximately 40 million in the year 2000 (Panescu, 2006). An aging population will necessitate the need for better implant designs in the future. Studying the failure mechanisms of implants will help researchers design better versions of existing prostheses. The US has the biggest market share for medical devices, with cardiovascular devices (e.g. stents, defibrillators) representing the major sector and orthopedic devices (e.g. artificial joint replacements and limb replacement devices) being second in sales (Panescu, 2006). These two segments of the medical device market represent a significant portion of implants being used currently. However, research and development of other prostheses, including breast implants, artificial skin and small joint replacements, need to be carried out for our future patients. Evidence based studies in different population groups need to be encouraged. Meta-analyses, systematic reviews of randomized controlled trials and cohort studies need to be conducted in order to better delineate implant survival among different types, as well as their treatment outcomes (Lee et al., 2000). Observational, prospective and retrospective studies of implants and biomaterials can provide useful information, including some factors that might be responsible for their failure (Buser et al., 1997; Roumanas et al., 2000; Dobbs, 1980; Roos et al., 1997; Brocard et al., 2000; Jones et al.,

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2007; Cook et al., 1999; Hatanaka et al., 2006; Liang et al., 2006). Doubleblind controlled studies are essential to provide medical and dental practitioners with Level I evidence as required by Evidenced-Based Medicine, which is presently considered the gold standard in determining the practice guidelines. A major benefit of such endeavors will be to use the results in large population segments worldwide. Consequently, reductions in the number of implant failures and savings in health care costs can be expected.

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khan ma, williams rl and williams df (1996). In-vitro corrosion and wear of titanium alloys in the biological environment. Biomaterials, 17, 2117–2126. doi: 10.1016/0142-9612(96)00029-4. kido h, schulz ee, kumar a, lozada j and saha s (1997). Implant diameter and bone density: Effect on initial stability and pull-out resistance. J Oral Impl, 23, 163–169. kim jw, kim jh, thompson vp and zhang y (2007). Sliding contact fatigue damage in layered ceramic structures. J Dent Res, 86, 1046–1050. doi: 10.1177/ 154405910708601105. korovessis p, petsinis g, repanti m and repantis t (2006). Metallosis after contemporary metal-on-metal total hip arthroplasty – Five to nine-year follow-up. J Bone Joint Surg-(Am), 88A, 1183–1191. lee jj, rouhfar l and beirne or (2000). Survival of hydroxyapatite-coated implants: A meta-analytic review. J Oral Maxillofac Surg, 58, 1372–1379. doi: 10.1053/ joms.2000.18269. liang b, fujibayashi s, fujita h, ise k, neo m and nakamura t (2006). Long-term follow-up study of bioactive bone cement in canine total hip arthroplasty. J Longterm Eff Med Impl, 16, 291. lin d, li q, ichim i and swain m (2007). Damage evaluation of bone tissues with dental implants. Key Eng Mater, 348, 905. doi: 10.4028/www.scientific.net/ KEM.348-349.905. lykissa e and maharaj s (2006). Platinum concentration and platinum oxidation states in body fluids, tissue, and explants from women exposed to silicone and saline breast implants. J Long-term Eff Med Impl, 16, 435. malchau h, herberts p and ahnfelt l (1993). Prognosis of total-hip replacement in Sweden – follow-up off 92 675 operations performed 1978–1990. Acta Orthop Scand, 64, 497–506. marshall a, ries md and paprosky w (2007). How prevalent are implant wear and osteolysis, and how has the scope of osteolysis changed since 2000? AAOS/NIH Osteolysis and Implant Wear: Biological, Biomedical Engineering, and Surgical Principles. Austin, TX. mccracken m, lemons je, rahematulla f, prince cw and feldman d (2000). Bone response to titanium alloy implants placed in diabetic rats. Int J Oral Maxillofac Impl, 15, 345–354. mercuri lg and anspach we (2003). Principles for the revision of total alloplastic TMJ prostheses. Int J Oral Maxillofac Surg, 32, 353–359. doi: 10.1016/ j.joms.2004.06.005. mercuri lg and giobbie-hurder a (2004). Long-term outcomes after total alloplastic temporomandibular joint reconstruction following exposure to failed materials. J Oral Maxillofac Surg, 62, 1088–1096. doi: 10.1016/j.joms.2003.10. 012. misch ce (1999). Implant design considerations for the posterior regions of the mouth. Impl Dent, 8, 376–86. mombelli a and cionca n (2006). Systemic diseases affecting osseointegration therapy. 1st Consensus Conference of the European-Association-for-Osseointegration. Pfaffikon, Switzerland. morris hf, ochi s and winkler s (2000). Implant survival in patients with Type 2 diabetes: Placement to 36 months. Ann Periodontol, 5, 157–165.

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moy pk, medina d, shetty v and aghaloo tl (2003). Dental implant failure rates and associated risk factors. 18th Annual Meeting of the Academy-of-Osseointegration. Boston, MA. nassiri n, cleary dr and ueeck ba (2009). Is cranial reconstruction with a hardtissue replacement patient-matched implant as safe as previously reported? A 3-year experience and review of the literature. J Oral Maxillofac Surg, 67, 323–327. doi: 10.1016/j.joms.2008.08.032. neut d, van horn jr, van kooten tg, van der mei hc and busscher hj (2003). Detection of biomaterial-associated infections in orthopaedic joint implants. Clin Orthop Rel Res, 261–268. okazaki y, rao s, ito y and tateishi t (1998). Corrosion resistance, mechanical properties, corrosion fatigue strength and cytocompatibility of new Ti alloys without Al and V. Biomaterials, 19, 1197–1215. doi: 10.1520/JAI12783. panescu d (2006). ‘Medical Device Industry’. Wiley Encyclopedia of Biomedical Engineering, 1–10, John Wiley & Sons, New York. pilliar rm, lee jm and maniatopoulos c (1986). Observations on the effect of movement on bone ingrowth into porous-surfaced implants. Clin Orthop Rel Res, 208, 108–113. prevention (2007). Diabetes Data & Trends, vs Center for Disease Control and Prevention (CDC). rahaman m, yao a, bal b, garino j and ries m (2007). Ceramics for prosthetic hip and knee joint replacement. J Am Ceram Soc, 90, 1965–1988. doi: 10.1111/j.1551-2916.2007.01725.x. ratner bd and hench l (1999). Perspectives on biomaterials. Curr Opin Sol St Mater Sci, 4, 379–380. doi: 10.1016/S1359-0286(99)00051–0. reclaru l and meyer jm (1994). Study of corrosion between a titanium implant and dental alloys. J Dent, 22, 159–168. restrepo c, parvizi j, kurtz sm, sharkey pf, hozack wj and rothman rh (2008). The noisy ceramic hip: Is component malpositioning the cause? J Arthrop, 23, 643–649. doi: 10.1016/j.arth.2008.04.001. revell pa (2008). The combined role of wear particles, macrophages and lymphocytes in the loosening of total joint prostheses. J Royal Soc Int., 5, 1263–1278. doi: 10.1098/rsif.2008.0142. revell pa, alsaffar n and kobayashi a (1997). Biological reaction to debris in relation to joint prostheses. Proc Instn Mech Engrs Part H-J Engng Med, 211, 187–197. doi: 10.1243/0954411971534304. roos j, sennerby l, lekholm u, jemt t, grondahl k and alberktsson t (1997). A qualitative and quantitative method for evaluating implant success: A 5-year retrospective analysis of the Branemark implant. Int J Oral Maxillofac Impl, 12, 504–514. rose rm, schiller al and radin el (1972). Corrosion-accelerated mechanical failure of a vitallium nail-plate. J Bone Joint Surg-(Am), 54A, 854–862. roumanas ed, freymiller eg, chang tl, aghaloo t and beumer j (2000). Implantretained prostheses for facial defects: An up to 14-year follow-up report on the survival rates of implants at UCLA. Joint Symposium of the American-Academyof-Maxillofacial-Prosthetics/International Congress of Maxillofacial Prosthetics. Kauai, Hawaii, Quintessence Publ Co Inc. roychowdhury a, pal s and saha s (2000). Stress analysis of an artificial temporal mandibular joint. Crit Rev Biomed Eng, 28, 411–420.

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schmalzried tp, peters pc, maurer bt, bragdon cr and harris wh (1996). Long duration metal-on-metal total hip arthroplasties with low wear of the articulating surfaces. J Arthrop, 11, 322–331. doi: 10.1016/S0883-5403(96)80085-4. shalabi mm, gortemaker v, vanthof ma, jansen ja and creugers nhj (2006). Implant surface roughness and bone healing: A systematic review. J Dent Res, 85, 496–500. doi: 10.1177/154405910608500603. sieber hp, rieker cb and kottig p (1999). Analysis of 118 second-generation metalon-metal retrieved hip implants. J Bone Joint Surg-(Br), 81B, 46–50. speculand b, hensher r and powell d (2000). Total prosthetic replacement of the TMJ: Experience with two systems 1988–1997. Br J Oral Maxillofac Surg, 38, 360–369. doi: 10.1054/bjom.2000.0338. stulberg bn, watson jt, stulberg sd, bauer tw and manley mt (1991). A new model to assess tibial fixation. II. Concurrent histologic and biomechanical observations. Clin Orthop Rel Res, 263, 303–309. sudhakar kv (2005). Investigation of failure mechanism in vitallium 2000 implant. Eng Fail Anal, 12, 257–262. doi: 10.1016/j.engfailanal.2004.05.005. sumner dr, kienapfel h, jacobs jj, urban rm, turner tm and galante jo (1995). Bone ingrowth and wear debris in well-fixed cementless porous-coated tibial components removed from patients. J Arthrop, 10, 157–167. doi: 10.1016/ S0883-5403(05)80122-6. sutow ej, jones dw and milne el (1985). In vitro crevice corrosion behavior of implant materials. J Dent Res, 64, 842–847. doi: 10.1177/00220345850640051201. sykaras n, iacopino am, marker va, triplett rg and woody rd (2000). Implant materials, designs, and surface topographies: Their effect on osseointegration. A literature review. Int J Oral Maxillofac Impl, 15, 675–690. ta le, phero jc, pillemer sr, hale-donze h, mccartney-francis n, kingman a, max mb, gordon sm, wahl sm and dionne ra (2002). Clinical evaluation of patients with temporomandibular joint implants. J Oral Maxillofac Surg, 60, 1389–1399. doi: 10.1053/joms.2002.36089. takeshita f, iyama s, ayukawa y, kido ma, murai k and suetsugu t (1997). The effects of diabetes on the interface between hydroxyapatite implants and bone in rat tibia. J Periodont, 68, 180–185. takeshita f, murai k, iyama s, ayukawa y and suetsugu t (1998). Uncontrolled diabetes hinders bone formation around titanium implants in rat tibiae. A light and fluorescence microscopy, and image processing study. J Periodont, 69, 314–320. taylor s, manley mt and sutton k (2007). The role of stripe wear in causing acoustic emissions from alumina ceramic-on-ceramic bearings. J Arthrop, 22, 47–51. triplett rg, frohberg u, sykaras n and woody rd (2003). Implant materials, design, and surface topographies: Their influence on osseointegration of dental implants. J Long-term Eff Med Impl, 13, 485–501. unwin ps, cannon sr, grimer rj, kemp hbs, sneath rs and walker ps (1996). Aseptic loosening in cemented custom-made prosthetic replacements for bone tumours of the lower limb. J Bone Joint Surg-(Br), 78B, 5–13. upadhyay d, panchal ma, dubey rs and srivastava vk (2006). Corrosion of alloys used in dentistry: A review. Mat Sci Eng A-Struct Mat Prop Micro Proc, 432, 1–11. urban rm, jacobs jj, gilbery jl and galante o (1994). Migration of corrosion products from modular hip prostheses – particle microanalysis and histopathological findings. J Bone Joint Surg-(Am), 76A, 1345–1359.

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vaderhobli rm and saha s (2007). Profilometric surface roughness analysis of Christensen metal temporomandibular joint prostheses. J Long-term Eff Med Impl, 17, 281–288. van de belt h, neut d, schenk w, van horn jr, van der mei hc and busscher hj (2001). Infection of orthopedic implants and the use of antibiotic-loaded bone cements – A review. Acta Orthop Scand, 72, 557–571. doi: 10.1080/ 000164701317268978. vinh d and embil j (2005). Device-related infections: A review. J Long-term Eff Med Impl, 15, 467. walter wl, lusty pj, watson a, o’toole g, tuke ma, zicat b and walter wk (2006). Stripe wear and squeaking in ceramic total hip bearings. Sem Arthrop, 17, 190–195. westermark a, koppel d and leiggener c (2006). Condylar replacement alone is not sufficient for prosthetic reconstruction of the temporomandibular joint. Int J Oral Maxillofac Surg, 35, 488–492. doi: 10.1016/j.ijom.2006.01.022. widmer a (2001). New developments in diagnosis and treatment of infection in orthopedic implants. Clin Infect Dis, 33, 94–106. doi: 10.1086/321863. wright tm, rimnac cm, stulberg sd, mintz l, tsao ak, klein rw and mccrae c. (1992). Wear of polyethylene in total joint replacements – observations from retrieved PCA knee implants. Clin Orthop Rel Res, 126–134. wright tm and goodman sb (2001). ‘What surgical-related factors contribute to implant wear?’ In Wright TM and Goodman SB (Ed.) Implant Wear in Total Joint Replacement. Rosemont, American Academy of Orthopedic Surgeons. yamaguchi m, bauer tw and hashimoto y (1997). Three-dimensional analysis of multiple wear vectors in retrieved acetabular cups. J Bone Joint Surg-(Am), 79A, 1539–1544. zhang ds, lu cl, zhang xy, mao ss and arola d (2008). Contact fracture of fullceramic crowns subjected to occlusal loads. J Biomech, 41, 2995–3001. doi: 10.1016/j.jbiomech.2008.07.019. zhang y and lawn b (2005). Competing damage modes in all-ceramic crowns: Fatigue and lifetime. Bioceramics, 17, 284–286, 697–700. zhang y, pajares a and lawn br (2004). Fatigue and damage tolerance of Y-TZP ceramics in layered biomechanical systems. J Biomed Mater Res Part B-App Biomat, 71B, 166–171. doi: 10.1002/jbm.b.30083.

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13 Rapid prototyping in biomedical engineering: structural intricacies of biological materials S. J. K A L I TA, University of North Dakota, USA

Abstract: Progress in tissue engineering promises patients new biomedical technologies to reduce pain, regenerate tissue, restore structure and functions, and on the horizon, replace failing organs with fully functional artificial ones. Applications of rapid prototyping (RP) in tissue engineering offer the pledge of growing these regenerative tissues and functional organs. RP has been effective in integrating structural architecture and assimilation of hormones in scaffolds. The chapter discusses several RP technologies and their applications in biomedical engineering in mimicking the structural intricacies of biological materials. A synopsis on biomaterials, materials properties of structural biomaterials and custom-designed scaffolds is also included. Key words: rapid prototyping, biomedical engineering, scaffold design, structural intricacies, hard tissues, laser additive manufacturing, solid free-form fabrication.

13.1

Introduction

The relatively recent maturation of materials science and its transition from an empirical discipline to a scientific one, has opened doors to many revolutionary accomplishments in engineering. The adoption of a quantitative scientific approach in materials science and engineering has led to the development of highly specialized materials that are able to perform unique tasks for the benefit of mankind. This is evident in the field of biomedical engineering, where two major disciplines merged to create new and practical ways to overcome age-old diseases and other challenges associated with health care. Due to the birth of biomedical engineering, which dates back to the turn of the 20th century, the success of long-term artificial implants and prostheses during the 1950s,1 and advanced nanomedicine in the 2000s, commonplace diseases and complicated conditions are treated effectively using real-world solutions. Across the world, musculoskeletal conditions alone affect hundreds of millions of people at a huge cost to society.1 More than half of all injuries are to the musculoskeletal system. Back or spine impairments, which number about 18.4 million, are the most prevalent musculoskeletal conditions for persons aged 18 and older. Musculoskeletal conditions are the 349 © Woodhead Publishing Limited, 2010

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major causes of severe long-term pain, physical disability and loss of workhours.1,2 In the United States alone, 28.6 million people incur a musculoskeletal injury every year; these conditions cost $254 billion each year.1 Medical procedures to address bone related injury are prevalent in the United States, with around 900 000 hospitalizations due to fractures and over 800 000 grafting procedures annually. Diseases such as osteoporosis, osteoarthritis, bone metastases and spinal impairments are major threats to human health. Osteoporosis is a disease characterized by low bone mass and structural deterioration of bone tissue, leading to bone fragility and fractures, especially of the hip, spine, and wrist. The skeleton is also a favored site for metastasis of human cancer.3 Breast and prostate cancer cells often metastasize to the bone,3,4 where they stimulate osteoclast osteolysis4 and cause systemic bone destruction.5 One out of every two women and one out of every eight men, older than age 50, will have an osteoporosis-related fracture in their lifetime. Spinal disorders lead to an overall compromise in the patient’s quality of life. It is estimated that the number of people affected by arthritis will increase to 60 million by 2020. Tissue grafting and the use of implants have become a routine procedure for orthopedic and spine surgeons in treating various musculoskeletal conditions. Tissue grafting has been known for a long time. There is archeological evidence that lost teeth were replaced by hand-carved ivory or wood ‘implants’ as early as ancient Egyptian times. Tissue grafting can be divided into four major categories according to the genetic relationship between the donor and the recipient. They are: (a) autogenous tissue graft (autograft): a tissue graft from one site to another within the same individual; (b) isogenous tissue graft (isograft): donor and recipient are genetically identical individuals of the same species (monozygotic twins, animals of highly inbred strain); (c) allogenous tissue graft (allograft): a tissue graft between individuals of the same species; (d) xenogenous tissue graft (xenograft): donor and recipient are individuals from different species. The alternative to tissue grafting is replacement with artificial, designed scaffolds or implants fabricated using various biomaterials. Autograft is considered the ‘gold standard’ in the biomedical industry. However, there are serious disadvantages associated with this technique, which includes donor site morbidity, limitations in mechanical strength, limited availability of tissue for grafting and the need to undergo two surgical operations. More serious problems, such as transmission of diseases, and rejection by the patient’s immune system triggered by an immunological response due to genetic differences, are associated with the other three bone grafting techniques. Paradoxically, the morbidity and pain associated with the harvesting procedure for autografting is often greater than the primary surgery. Elimination of the second surgery also reduces hospitaliza-

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tion stay, complications associated with it, and costs. For these reasons, there is a rationale for designing and developing artificial scaffolds for tissue engineering applications and the demand is growing steadily. This has been a recurrent theme for the implant industry. Swedish orthopedic surgeon Per Ingvar Branemark studied the healing process of titanium screws in bone repair, during the early 1950s. He observed osseointegration – fusion of titanium screws with bone.6 His work generated momentum for the development of structural implants for clinical applications. Tissue engineering is a multidisciplinary field which involves the ‘application of the principles and methods of engineering and life sciences towards the fundamental understanding of structure–function relationships in normal and pathological mammalian tissues and the development of biological substitutes that restore, maintain or improve tissue function’.7 Tissue engineering and regenerative medicine is an evolving interdisciplinary field that integrates aspects of engineering and other quantitative sciences with biology and medicine, for the development of functional tissues and organs in vitro, for implantation in vivo or for direct remodeling and regeneration of tissue in vivo to repair, replace, preserve or enhance tissue or organ function lost due to disease, injury, or aging. The goal of tissue engineering is to surpass the limitations of conventional treatments based on organ transplantation and biomaterial implantation.8 Although, it is the vision of tissue engineering to create threedimensional functional tissues and organs, synthetically, cells lack the ability to grow three dimensionally into the desired anatomical shapes of tissues and organs. Instead, in vitro culture of cells grows in two-dimensional layers. In order to achieve a three-dimensional structure, it is imperative to seed and grow cells onto structures (preferably porous), known as ‘scaffolds’. Therefore, the design and fabrication of scaffolds with required properties is a key component for tissue engineering. The design and development of scaffolds with optimum properties for specific applications present serious challenges. In addition to the issues related to materials, processing and manufacturing, one must evaluate long-term biocompatibility of the scaffolds. Also, designing scaffolds to mimic structural intricacies of the tissue being replaced, customized for each patient, is practically impossible through traditional manufacturing approaches. Nevertheless, several advanced manufacturing/fabricating technologies developed during the last two decades, could overcome the limitations of traditional manufacturing technologies in creating complex scaffolds for tissue engineering applications. One of those processes is rapid prototyping (RP), which encompasses a range of different techniques based on fundamental science and engineering. RP normally refers to techniques that have the ability to fabricate physical objects automatically through continual build-up or creation of solid

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material through additive manufacturing. Consequently, RP is fundamentally different from conventional forming and machining types of manufacturing, and has unique capabilities for constructing structures with architectural intricacies. RP processes have shown significant potential for developing patient-specific scaffolds with different structural properties using several kinds of materials. RP has been effective in integrating structural architecture with changes in surface chemistry of the scaffolds, and assimilation of growth hormones. In addition, revitalization of imaging techniques has transformed diagnostics, scaffold and implant technology, and surgical procedures. The evolutionary advances in microelectronics, sensor technology, biomaterials engineering, RP technologies, design and analysis software, and advanced MEMS and NEMS devices, will help create novel scaffolds, implants and therapeutic devices with improved functionality and longevity.

13.1.1 Porous scaffolds in bone repair – a brief review A porous scaffold with open porosity resembles the structure of a cancellous or spongy bone. In the natural process, cancellous bone is converted to compact bone through progressive deposition of new bone lamellae by a series of osteons. Various processing techniques have been utilized to fabricate porous scaffolds to enhance tissue regeneration and bone repair. Most of these conventional processes had been used to create porous ceramic scaffolds. Many of these conventional scaffold fabrication techniques could create scaffolds having a continuous, uninterrupted pore structure with varying pore sizes which lacks any long-range channeling or three-dimensional interconnectivity. The replamineform process (one commercially successful technique) has been utilized to fabricate inert, bioactive, ceramic and polymeric implants that duplicate the macroporous microstructures of the corals which have interconnected pores. In this approach, natural corals are first machined to the desired shapes. The machined coral is fired to drive-off the carbon dioxide from the limestone (CaCO3) leaving behind a calcia structure to be used as an investment material to form the porous structure. The desired material is then cast into the pores and the calcia mold is removed from the structure by dissolving in dilute hydrochloric acid. Genus Porites is used to mimic the osteon-evacuated stroma of cortical bone, whereas to mimic cancellous bone, genus Goniopora is used. Porous and shaped hydroxyapatite (HA) implants, fabricated using the replamineform process, have been used for craniofacial reconstruction and have shown rapid bone in-growth. Their corresponding HA forms are known as HA200 and HA500, respectively. Dense hydroxyapatite and other biocompatible ceramic scaffolds show little in-growth of bone cells in animals. Although the replamineform

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process allows the fabrication of porous ceramics with pore sizes varying from 140–1000 μm, there is almost no control over the pore sizes, their distribution or their interconnectivity. Researchers have also developed porous ceramic scaffolds for tissue engineering using pore formers. Porous ceramic structures can be fabricated using ceramic slurry mixed with foaming agent followed by sintering at elevated temperature. The drawback of this technique is that there is also almost no control over pore size distribution and interconnectivity in the final structure. In spite of all the attempts to fabricate porous bioactive ceramic scaffolds for bone repair; to date, bioactive ceramics are most widely used either as powders or in particulate form to fill bone defects. Porous bioactive ceramic scaffolds lack the necessary mechanical strength to withstand the complex states of stresses experienced in various anatomical structures. Similar approaches have been used to create porous scaffolds Table 13.1 Conventional manufacturing techniques used to create porous polymeric scaffolds for tissue engineering Technique

Author and year

Scaffold material

Emulsion freeze drying

Whang et al. 19959

PGA

10

Fiber meshes/fiber bonding

Cima et al. 1991 Mikos et al. 199311 Freed et al. 199412

PLLA, PGA PGA PGA

Freeze drying

Hsu et al. 199713 Yannas et al. 198014 Doillon et al. 198615, Schoof et al. 200116 Madihally et al. 199917 Glicklis et al. 200018

PLGA Collagen 3D pore structure

Gas foaming

Mooney et al. 199619

PLGA, 100–500 μm pore size

Melt molding

Thompson et al. 199520

PLGA, HA

Phase separation

21

Lo et al. 1995

Guan et al. 200522 Guan et al. 200723

Chitin Alginate

Biodegradable synthetic polymer Polyurethane Elastomeric poly(ester urethane) urea (PEUU)

Solution casting

Reuber et al. 198724 Okuno et al. 199325

Polyurethane PVC membrane

Solvent casting particulate leaching

Mikos et al. 199426, Mikos et al. 199627

PLLA

PGA, polyglycolide; PLLA, poly(L-lactic acid); HA, hydroxyapatite; PVC, polyvinylchloride.

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using polymeric and polymer–ceramic composite materials. Table 13.1 lists some of the conventional processing techniques used by researchers to create porous polymeric scaffolds for tissue engineering. In recent years, RP and solid freeform fabrication (SFF) techniques have been successfully explored to fabricate controlled porosity ceramic scaffolds with varying pore size, pore shape, pore volume and three-dimensional interconnectivity. RP and SFF are common names for a group of additive manufacturing processes that use a layered manufacturing approach to fabricate 3D objects on a fixtureless platform, directly from computer-aided design data in a layer-by-layer manner. RP and SFF techniques offer significant advantages over conventional processes of creating scaffolds for tissue repair, and therefore hold a promising future. This chapter discusses the application of various RP techniques in scaffolding technology for tissue engineering from the perspective of complexity and challenges associated with mimicking structural intricacies and architectural features of natural tissues.

13.2

An overview of biomaterials

Developments in hard tissue engineering are directly attributed to innovations and discoveries in materials science and technology. Challenges in tissue engineering have motivated scientists and engineers to develop newer biomaterials, with improved properties, that can restore the structural features and physiological functions of natural tissues without adversely damaging the nearby tissues. The term ‘Biomaterials’ can be effectively defined as ‘materials of synthetic as well as of natural origin in contact with tissue, blood, and biological fluids, and intended for use for prosthetic, diagnostic, therapeutic, and storage applications that do not adversely affect the living organism and its components’.28 Over the last half century, applications of novel materials for biomedical practices have greatly revolutionized the quality of human life. Many specialty polymers, metals, ceramics and composites have been developed for numerous applications. These materials, in various forms and phases, are made to perform various functions in repair and reconstruction of diseased and damaged parts of the human body. Nevertheless, the successful in vivo applications of biomaterials depend on several different factors, such as the physical and mechanical properties, design, cytotoxicity, biocompatibility, bioactivity and biodegradability of the material used, as well as other factors which are not under the control of engineers/manufacturers, including the surgical techniques used by the surgeons, general health and condition of the patient, and post-surgery activities of the patient. A large number of materials (metals, ceramics, polymers and composites) of natural and synthetic origin have been developed or used in making

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scaffolds, to restore lost structure and functions of diseased or damaged tissues in clinics. The potential for use of any material as a biomaterial relies upon its compatibility with the physiological environment. Evolution of specially designed polymers, such as polylactide (PLA), polyglycolide (PGA), polyhydroxybutyrate (PHB) and their copolymers, has rejuvenated the field of tissue engineering from drug delivery to bone augmentation. Based on tissue response at the interface, biomaterials used in bone tissue engineering can be broadly classified into three categories. They are listed below. 1

Bioinert materials: a bioinert material forms a non-adherent fibrous capsule that ‘walls-off’ or isolates the implants from the host, which leads to complete encapsulation of an implant within the fibrous layer. This results in long-term loosening of implants.29 Examples include all metallic biomaterials, alumina, and polyethylene. 2 Bioactive materials: a bioactive material elicits a specific biological response at the interface of the material, which results in the formation of a bond between the tissues and the material. Examples include hydroxyapatite, polylactides and bioactive glass ceramics. 3 Bioresorbable materials: a bioresorbable material dissolves in the body and produces by-products that can be absorbed/digested by the body through various metabolic activities. These are ideal for many tissue engineering applications.29 Examples include tricalcium phosphates and polyethylene glycols. Metallic biomaterials are used in implant fabrication mainly because of their excellent mechanical properties and the extensive knowledge-base of mankind with regard to their processing, properties and structures. They are mostly used as passive substitutes and as supports for hard tissue repair and/or replacement in load-bearing applications such as total hip and knee joint replacements, for fracture healing aids such as bone plates, screws and wires, spinal fixation devices, and for dental implants, because of their excellent mechanical properties and good resistance to corrosion and wear. Some metallic biomaterials are used for more active roles in devices such as vascular stents, catheter guide wires, orthodontic arch wires and cochlea implants. Problems associated with metallic implants include stress shielding, corrosion, wear, and loosening of implants. The thermal and chemical stability of ceramics, high strength, wear resistance and durability, all contribute to making ceramics good candidate materials for surgical implants. Bioceramics, especially calcium phosphates, are preferred materials as bone grafts because of their excellent chemical stability, low density and compositional similarities with bone mineral. Though the potential uses of bioceramics could be massive, the field is still budding, with very little progress. It is true that we are yet to master bioceramic systems in depth. High strength bioceramics such as alumina and

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zirconia are also in use for load-bearing applications (e.g. ball and socket of total hip replacement). But like metals, these ceramics are bioinert, and form non-adherent fibrous capsules in vivo, leading to loosening of implants. In contrast, bioactive ceramics interact with the surrounding tissues and form strong interfacial bonds. They are composed of ions commonly found in physiological environments, which make them highly biocompatible. Moreover, they are resistant to microbial attack, pH changes and are stable during temperature changes. Unfortunately, these ceramics exhibit poor mechanical strength and low crack growth resistance, which limit their uses to non-load-bearing applications. In addition, the bioactivity of synthetic calcium phosphates is inferior to that of bone mineral. Tissue engineering has provided an alternative medical therapy using implants of polymeric biomaterials, with or without living precursor cells, as opposed to various transplants. A number of different polymers can be used as scaffolds to promote cell adhesion and maintenance of differentiated cell functions without hindering proliferation. Synthetic and natural polymers can also be used as templates to organize and direct the growth of cells and to help in the function of extracellular matrices (ECM).30 Both natural and synthetic polymers have become a part of regular, as well as advanced, medical care in the biomedical industry. The most commonly investigated natural polymers include alginate, collagen, hyaluronic acid and chitosan. Much interest in these natural polymers is because of their biocompatibility, relative abundance, ease of processing and/or their possible ability to imitate the microenvironment found within cartilage. Synthetic polymeric materials have been widely used in medical disposable supplies, prosthetic and dental materials, implants, dressings, extracorporeal devices, encapsulants, polymeric drug delivery systems, tissue engineered products, and in orthopedics. The main advantages of polymeric biomaterials compared to metal and ceramic materials are ease of manufacturability to produce various shapes, ease of secondary processibility, reasonable cost, and availability with desired mechanical and physical properties. Composite materials offer a variety of advantages over metals, ceramics and polymers, as they can incorporate the desirable properties of each of the constituent materials, while mitigating the more limited characteristics of each component. Generally, a composite can be defined as a materials system composed of a mixture or combination of two or more constituents that differ in form and chemical composition and are essentially insoluble in each other. The properties of composite materials depend upon the shape of the heterogeneities, upon the volume fraction occupied by them, and upon the interfaces among the constituents. Most of the biological materials found in nature are composites; for example, bone, skin, wood and carti-

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lages. The main benefit of using composite biomaterials is the ability to tailor their properties as per need and thus provide significant advantage over homogeneous biomaterials. Composite biomaterials can be classified as either bioinert, bioactive or bioresorbable. Some current applications of composite biomaterials in the biomedical industry include dental composites used as filler materials, coated metallic implants and reinforced polymethyl methacryalate (PMMA) bone cements used in many joint replacement surgeries. All kinds of composite biomaterials with different matrices, e.g. metallic, ceramic and polymeric, have been developed and tested for biomedical uses. Among them, greater emphasis has been given to polymer–ceramic composite biomaterials in recent years.

13.3

Material properties of structural biomaterials

It has been well established that the success of biomaterials requires the simultaneous achievement of a stable interface with connective tissues and a match of the mechanical behavior of the implant with the tissue to be replaced. Therefore, properties of materials play a vital role in short-term acceptability and long-term survivability of scaffolds and implants. Important properties of materials used in scaffold technology in hard tissue engineering are their physical, mechanical, surface and biological properties.

13.3.1 Physical and chemical properties The physical and chemical properties of implants are very important for processing and performance, as they are directly related to the mechanical and biological properties of the implants. Physical properties of materials, such as apparent density, absolute density, porosity, microstructure, crystal structure and degree of crystallinity, specific heat, melting point and boiling point, are directly related to mechanical properties such as tensile strength, bending strength, compressive strength and hardness, and they also determine the processing conditions or parameters. Bulk material properties are dependent on inter-atomic bonding (ionic, metallic or covalent), while the surface properties of materials are governed by physical forces such as van der Waals attraction and electrosteric repulsion. Although synthetic biomaterials and scaffolds have a wide range of promising applications in regenerative medicine and tissue engineering, their applications are limited by the basic materials’ properties. Other unique properties of materials, such as shape memory behavior, have been helpful for their use in biomedical applications. Shape memory alloys (SMAs) have rendered themselves favorable for use in the medical field primarily because of improved mechanical, chemical and biological

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properties, which include the capability of resistance against in vivo degradation, decomposition, corrosion, etc. Some of the more prominent medical applications include orthopedics (alloy implantation in the body for long periods), dentistry, and components in medical devices and instruments. One of the first biomedical applications that achieved broad acceptance was the superelastic SMA for dental arch wires. Shape-memory bone fasteners are proving superior to the usual screwed metal plates when it comes to joining broken bones. They are also widely used in making stents.

13.3.2 Mechanical properties Evaluation of mechanical properties of materials is important for matching biomaterial properties to the in vivo microenvironment. Mechanical property requirements are tied to specific applications. Different components of an implant generally have different mechanical property requirements. For example, the stem of a total hip replacement (THR) requires good tensile strength, bending and fatigue strength, while the ball and the socket of a THR (acetabular components) are required to have excellent wear resistance and resistance to impact. Accordingly, the stem of THR is made of Ti-6Al-4V alloy or a specific Co-Cr alloy for their strength, while alumina or zirconia is the material of choice for the ball in current technology due to their very high wear resistance. On the other hand ultra high molecular weight polyethylene (UHMWPE) sockets are gaining more importance as they can absorb greater impact. However, other materials like Co-Cr alloy are also used to make the ball and socket because of very high wear resistance. In polymers, the poor mechanical strength of PLLA, PGA and their copolymers has still been a bottleneck in creating load-bearing implants, although these materials are well known for their excellent bioactivity and bioresorbability. Dynamic mechanical tests are generally performed to determine time-dependent viscoelastic behavior of polymeric materials. The glass transition temperature and viscoelastic properties are of very high concern in polymer processing. For devices made of degradable materials, strength loss as a function of time due to biodegradation is an important consideration. Nevertheless, mechanical properties such as Young’s modulus, specific modulus, tensile strength, compressive strength, shear strength, yield strength, ductility and Poisson’s ratio are critical to evaluate before scaffold design and fabrication.

13.3.3 Surface properties Surfaces are critical to the study of biomaterials and the development of scaffolds or implants. The nature of an implant’s surface determines its

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interaction with the body fluids, which contains several different kinds of proteins. Interaction of these proteins with biomaterial/implant surfaces leads to cascades of reactions comprising the body’s response to the implant and determining the development of the tissue and implant interface and long-term survivability. The surface properties of implants are therefore particularly important in order to address biocompatibility issues and the development of designs. This is not an easy task, in view of many different surface properties that are considered to play an important role in the reaction of the host body to the artificial material. The ideal biomaterial should have good mechanical properties, while the surfaces should have good biocompatibility, i.e. the material must possess not only suitable mechanical properties in order to function properly inside the human body, but must not be harmful for the host tissue as well. Implant surfaces should not induce acute inflammatory response. Since it is the surface of a material that first comes into contact with the biological surroundings, the surface properties are in charge of the initial rejection or acceptance of a foreign device. The rationale for surface modification is therefore straightforward: to retain the key physical and bulk properties of a biomaterial intact while modifying only the outermost surface, in order to tailor the biological properties in most implants. Because it can be very difficult to produce materials that possess the necessary bulk properties with a biocompatible surface, a material’s surface can be modified in such a way so as to enhance its bioactivity, without sacrificing its bulk material properties. This has been done with every kind of biomaterial, to optimize the body’s response to each given implant. For example, polymers used with blood have been treated to reduce thrombic response.31 Other polymers used in bone tissue engineering have been treated to enhance osteoblast adhesion and proliferation.32,33 Various metallic biomaterials are treated to be more biomimetic and thereby integrate more readily with the tissues they contact.34,35 Ceramic biomaterials are also treated, as in the case of alumina, which, when treated with heparin, becomes much more biocompatible with neural tissue.36 Many methods have been explored to modify the surfaces of biomaterials. All of these include some alteration of the chemical or physical characteristics of the surface; and can be accomplished by various means, such as ion beam processing, plasma surface modification, biochemical immobilization of biological molecules on the material surface, or direct alteration methods such as etching32 or introduction of pores37 during fabrication of the material. Modifying the chemical character of the surface often involves changing its hydrophobicity.32 Another important consideration is the problem of biofilms.38 There is no such thing as a truly sterile implant. When a material is placed in any body, it lacks the symbiotic

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bacteria present on all tissues that protect the host from infection by foreign bacteria. This lack causes an uncontrolled growth of harmful bacteria and results in potentially life-threatening infections.39 In order to circumvent this, surface modification of implant materials with antibiotic coatings40 or with coatings that prevent adhesion of bacteria41 have been developed.

13.3.4 In vivo biological properties Biological responses to an implant are affected by several factors, which include the physiological properties of the surface and the bulk, as well as adsorption of biological moieties and the immunological response of the system. It has always been a challenge for scientists to define the term biocompatibility or biocompatible, as both these expressions are qualitative and a material which is compatible for a particular application may not function satisfactorily at a different location in the body. It has been rightly stated by Black that the label ‘biocompatible’ suggests that the material described exhibits universally ‘good’ or harmonious behavior in contact with living tissue and body fluids.42 He also has provided a generic definition of ‘biocompatible’ (or ‘biocompatibility’) as ‘biological performance in a specific application that is judged suitable to that situation’.42 Today, biocompatibility is succinctly defined as ‘the ability of a material to perform with an appropriate host response in a specific situation’. An understanding of protein interactions with biomaterials, which is governed by the thermodynamic principles of protein adsorption, provides insight into cell biomaterial interactions. In vivo biological properties are also dependent on the properties and components of the extracellular matrix, cell adhesion, spreading and migration, cell cycle, cell differentiation and cell phenotype, while for blood interfacing implants, thrombosis and thrombo-embolism are some of the important and critical concepts that require studies on the extrinsic and intrinsic coagulation cascade and the role of platelets.

13.4

Rapid prototyping – a novel manufacturing approach

Since the 1980s RP or SFF or rapid manufacturing (RM) or additive manufacturing (AM) technologies have emerged as revolutionary manufacturing processes with inherent capability to rapidly fabricate parts in virtually any shape. RP or SFF or RM or AM, combined with computer-aided design (CAD) and computer-aided manufacturing (CAM), has the distinctive benefit of being able to build objects with predefined microstructure and macrostructure.43–46 This distinct advantage gives RP the potential for making scaffolds or orthopedic implants with controlled hierarchical

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structures.46–51 In addition, RP or SFF processes have the significant potential to drastically reduce the time lag between conceptual design of a structure and its actual fabrication. RP has proved to be a time-saving and economical process as compared to conventional manufacturing processes. RP takes virtual designs from computer design software, transforms them into thin, virtual, horizontal cross-sections and then creates successive layers until the model is complete. It is a WYSIWYG (what-you-see-is-what-you-get) process where the virtual model and the physical model are almost identical. Applications of RP technologies are wide ranging, including automobile and aerospace manufacturing, civil engineering, biomedical engineering, defense projects, consumer products and entertainment. RP instruments manufactured by Stratasys are used to make production parts for diesel fuel motorcycles and spacecraft apparatus. To make models, fused deposition modeling (FDM) is used to cast urethane prototypes in the automobile industry. Also, it is used to make automotive controls for the car. In the entertainment industry, RP is used to make 3D models for Hollywood and recently has been used for movies such as Jurassic Park-III. Nowadays, RP had found many applications in the medical field. For example, spinal implants are made using RP techniques; medical devices are made using stereolithography; bone implants made up of ceramic material and internal structures of bones to validate computer simulations. Moreover, doctors use RP to make virtual models during the planning of surgery; they also use biomodels to plan complicated procedures like separating Siamese twins, and in developing patient-specific dental implants. Several types of RP systems have been introduced into the market over the past 20 years and are being routinely used to manufacture parts for various applications. These RP systems can be classified in many different ways. Many of them are laser-based and are classified or named accordingly. Some of them use a powder bed while some others use a stream of powder. Some RP systems use filaments while there are others that use dense solutions. Depending on the process, RP products have different properties and advantages, as well as limitations. RP systems have also been developed using electron beams. Kruth et al. (1998) provided an early review of progress in additive manufacturing and RP technologies.52 Kruth et al. (1999) published another early review of laser-based RP technologies during its first decade from birth.53 Laser-based SFF processes use the heat generated by a focused beam of laser to partly or entirely melt the deposited material. The term ‘laser’ is an acronym for light amplification by stimulated emission of radiation. Lasers are typically spatially coherent, which means that the light is emitted either in a narrow, low-divergence beam, or can be converted into one with the help of optical components such as lenses. Table 13.2 presents a historical

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Biointegration of medical implant materials Table 13.2 Historical perspective on the development of laser-based rapid prototyping/SFF techniques Year

Inventor

Process

1910 1916 1940 1958 1960 1960 1974 1979 1985 1991 1990+

Schoop Einstein Brennan Schawlow & Townes Singer Mainman Osprey Breinan Hull Deckard –

Spray metallization Theory of light emission Spray metal strips Invented LASER/MASER Spray rolling, CSD, SSP Invented working Ruby Laser Spray forming LASERGLAZETM Stereolithography (STL) Selective laser sintering (SLS) DLFTM, DMDTM, LENS®, LAMTM

perspective on the development of laser-based RP/SFF techniques. Research efforts have demonstrated the great potential available within laser-based SFF to manufacture metal components with microstructures and mechanical properties equivalent or superior to conventionally processed materials (e.g. casting, hot isostatic pressing).54 During the 1990s, a large number of laser-based SFF processes were developed and commercialized for advanced manufacturing and processing of metal, ceramic, polymer and composite materials. Some of these techniques include laser-engineered net shaping (LENSTM), direct light fabrication (DLF), direct metal laser sintering (DMLS) and selective laser melting (SLM). These laser-based RP processes are capable of completely melting material from a powder bed or deposited as a stream of powder along with the laser. Movement of the laser with respect to the powder bed or simultaneous movement of powder with the laser on a substrate or bed creates the desired structures, coating or cladding. Most RP systems in the market primarily create external physical models and have the basic but critical limitation of one material on one stage. To overcome the limitations of mono-material and mono-function of conventional RP systems, Im et al. (2007) proposed the concept of functional prototype development (FPD).55 FPD provides the necessary prototype functions, such as mechanical, optical, chemical and electrical properties, in order to meet the broad requirements of the industry.55 Their experimental results demonstrated that FPD has great potential applied to broad industrial uses and that it will be a powerful tool in future.55 A brief overview of a few commonly used RP technologies is presented in the following sections.

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13.4.1 Laser additive manufacturing (LAM) or laser-based rapid prototyping LAM includes a number of processes that fundamentally use lasers to create structure through the addition of new material/s on a substrate or bed. Depending on the application, the actual structure may be removed from the bed. LAM offers several unique advantages, which include economic repair and remanufacturing, low heat input resulting in a minimal heat affected zone, minimal internal stress and distortion, strong metallurgical bonding, rapid cooling rate, very fast processing, improved wear and corrosion resistance, improved functionality, and technology versatility. Integration of 3D design software has enabled most LAM processes to fabricate structures in correct dimensions and complex geometries from CAD data descriptions. This has been a big help particularly in the field of biomedical engineering to design and develop scaffolds using homogeneous, as well as heterogeneous materials. LENSTM and DLF processes LENSTM technology, developed by Sandia National Laboratories, uses metal powders to create three-dimensional structures with functionality and improved strength as per a CAD model. The process uses up to 2 kW of Nd:YAG laser power focused onto a metal substrate to create a molten pool on the surface of the substrate. Metal powder is supplied coaxially to the focus of the laser beam through a deposition head. The bed or table has X–Y movement capable of moving in a raster fashion, while the head moves up vertically as each layer is built. An inert shielding gas is often used to eliminate oxidation and promote layer to layer adhesion through superior surface wetting. LENS is an interesting process for orthopedic device manufacturing, as it can directly build functional implants unlike other RP processes. Similar to LENSTM, the DLF process also melts powder that is injected into the laser beam. One of the limitations of these two processes is that they always require a solid substrate, and material wastage can be high due to powder scattering during processing. The process also involves complexities and challenges associated with powder flow rate, although both processes allow recycling and reuse of unmelted powder. Selective laser sintering (SLS) SLS is one of the earliest RP techniques, wherein a thin layer of powder is first deposited on the prototype/part build-area. The laser beam, guided by galvano mirrors, is scanned onto the powder bed, based on a CAD design

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strategy to form solidified or sintered layers. The surrounding powder acts as support and can be reused later, after screening. It is a layer-by-layer process and therefore the process continues until the part is completed. The typical thickness of each layer is 0.02 to 0.1 mm. The processing chamber is isolated and is generally filled with nitrogen or argon to avoid oxidation. Selective laser melting (SLM) or direct metal laser sintering (DMLS) SLS (late 1980) is a layer manufacturing/RP process. The success of SLS gave birth to the selective laser melting (SLM) process for creating functional high density metallic prototypes and parts. SLM is very similar to SLS in terms of equipment but uses a much higher energy density, which enables full powder melting and accordingly fabrication of structures having densities very close to the theoretical density. In SLM, near full-density parts are produced without the need for postprocessing. Powder is fully molten during processing. This results in thermal stresses, distortation, delamination and cracking in products. SLM is driven by the need to process near full-density objects with mechanical properties comparable to those of bulk materials.56–58 SLM uses a high power Nd:YAG laser to melt thin layers of metal powder. This makes processing of metals with high atomic number or high reflectivity easier because, with Nd:YAG lasers the energy absorbance is much higher.58 SLM or DMLS builds layers approximately 30 μm thick from atomized powder. The smaller size powder for SLM facilitates laser beam melting in a nitrogen environment with high energy density comparable to that of the EBM system (Section 13.5.3) in the build area. Unlike SLS, SLM is more difficult to control.57 Balling effect is a severe impediment to interlayer connection.57 Vaporization effect is noticed in SLM when the powder bed is irradiated with high energy intensities.56 SLM can create internal hollow structures and there are no geometric limitations, unlike conventional processes.

13.4.2 Fused deposition modeling (FDM) The fused deposition process, commercialized by Stratasys as fused deposition modeling (FDM) is a RP technique where three-dimensional (3D) objects are built layer-by-layer from a computer-aided design (CAD) file on a computer-controlled fixtureless platform.45,59 In this process, a thermoplastic polymer filament passes through a heated liquifier, where the liquifier is heated to a temperature that is slightly above the melting point of the build polymer. The liquifier extrudes a continuous bead or road of material through a nozzle and deposits it onto a fixtureless platform. The liquifier movement is computer controlled along the x and y directions,

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Filaments Rollers Support filament supply

X and Y stage system Extrusion nozzle Liquifier (heater)

Substrate

Z stage system

Object being fabricated

Model filament supply

Thermal housing (oven)

13.1 Schematic of fused deposition modeling process.

based on the build strategy of the part to be manufactured. When deposition of the first layer is complete, the fixtureless platform indexes down, and the second layer is built on top of the first layer. This process continues until the part-manufacturing is complete. The temperature of the liquifier and surrounding environment, as well as the filament feed rate and nozzle diameter, are some of the important variables that determine the quality of the final part. Figure 13.1 shows a schematic of the basic FDM process. The FDM process can be used in two different ways to fabricate porous ceramic structures, namely the direct FDM and the indirect FDM. Figure 13.2 shows a schematic of the customized direct FDM process showing the fabrication of a controlled porosity polypropylene–tricalcium phosphate (PP–TCP) composite scaffold with 3D interconnectivity. Figures 13.3a and b show the schematic of mold design illustrating road gap, road width and slice thickness, and the indirect FDM process. In the indirect FDM, porous polymeric molds are first fabricated via fused deposition using wax filaments. The molds are then infiltrated with ceramic slurry followed by binder removal and sintering to produce porous ceramic scaffolds. In the direct route, ceramic powders are mixed with an appropriate binder system and extruded into a continuous filament, which is then used as feedstock material to fabricate porous structures. These parts require post processing, such as binder removal and sintering. In addition to biomedical engineering applications, FDM is being used to make optically clear face shields for breathing apparatus. It has many

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Top view

PP–TCP composite filament

Top view Spool of filaments

X–Y direction liquifier

Porous PP–TCP scaffold

Prototype design

Z-direction platform with foam base

13.2 Schematic of FDM process, showing fabrication of a controlled porosity PP–TCP composite scaffold with 3D interconnectivity. PP–TCP composite filaments were developed using a PolyLab system torque rheometer from HaakeTM. These continuous filaments were used as feed stock material for FDM. (Reprinted from Materials Science and Engineering C, 23, 611–620, S.J. Kalita, S. Bose, H.L. Hosick, and A. Bandyopadhyay, ‘Development of controlled porosity polymer– ceramic composite scaffolds via fused deposition modeling’, Copyright (2003) with permission from Elsevier.68)

applications in manufacturing such as a replacement for casting because of its unique properties, also saving time and money. It is being used to make customized houses, building designs and even large objects such as Greek statues. Also, it is being used in building prototype airplanes for wind tunnel testing.

13.4.3 Electron beam melting (EBM) Electron beam melting (EBM) is a RP process, developed and commercialized by Arcam AB in Sweden; it produces fully dense metal parts directly from metal powder, having the characteristic properties of the target material. The EBM system builds structures from the bottom up by scanning the focused electron beam to selectively melt specific powder areas. It reads data from a 3D CAD model and lays down successive layers of powdered

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Y–spacing

(a) Raster

Double Z– spacing

X–spacing Z

Z

X (b)

CAD model

Y Y or X Porous polymer mold

Slurry infiltration

Sintered porous ceramic

13.3 Schematic showing (a) mold design development used in many indirect RP systems including FDM and (b) the indirect FDM process involving mold infiltration with slurry followed by sintering to create 3D porous ceramics as scaffolds for tissue repair.

material. The process continues until the last layer of the part is built. It takes place under vacuum, which makes it a highly suitable process to fabricate structures using reactive materials that cannot be exposed to the atmosphere. The EBM system is an electron optical system similar to an SEM60 or an EB melting system. The EBM process can build a minimum layer thickness of 0.05 mm and it has a tolerance capability of ±0.4 mm.

13.4.4 Three-dimensional printing (3DP) 3D printing is a new type of RP or additive manufacturing technology, where a 3D object is created using layer-by-layer manufacturing technology. In the 3DP process, a stream of adhesive drops is expelled through an inkjet print head, selectively bonding a thin layer of powder particles to form a solid shape.61 The resolution achieved is approximately 300 μm. 3D printers

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are generally faster, more affordable and easier to use than other additive manufacturing technologies. 3D printers offer potential for production of actual parts although the current market is dominated by prototypes for various applications. Some applications in scaffold technology have been explored with success. One of the variations of 3D printing is the inkjet printing system, which has been extensively used during recent years to create or print organs directly from CAD models. The TheriForm system is very similar to the 3DP system in operation. In the TheriForm system, a print head assembly deposits binder drops onto a chosen area of the powder bed, swelling and dissolving the polymer powder in the printed regions.

13.4.5 Three-dimensional Bioplotter (3D Bioplotter) The patented Bioplotter principle is simple. It is based on dispensing a plotting material into a plotting medium to cause solidification of the material and to compensate gravity force through buoyancy. The plotting material leaves the nozzle and solidifies in the plotting medium, subsequent to bonding to the previous layer. In general, when dispensing materials in semi-solid form, gravity causes complex structures to collapse. With the Bioplotter technique, the material is dispensed into a matched plotting medium, which compensates for the gravity through buoyancy.62

13.4.6 Three-dimensional fiber deposition (3DFD) The 3D fiber deposition RP process is very similar to the FDM process. In this technique, the feedstock material, in a pellet or granule form, is poured into a heated liquifier and, based on a predefined deposition strategy, the structure is built. The flow of the feedstock material is governed by applying pressure to the syringe containing it.

13.4.7 Pressure-assisted microsyringe (PAM) Pressure-assisted microsyringe (PAM) is a versatile and flexible system that can construct 2D and 3D high-resolution complex structures. PAM is based on the use of a microsyringe that utilizes a computer-controlled three-axis micropositioner. The microsyringe expels dissolved polymer (polymer dissolved in specific solvents) under a low and constant pressure to form the desired structure. The resolution of the deposition system is governed by the dimensions of the syringe tip. Often, high pressure is necessary to squeeze material from the small orifice. Precision polymer scaffolds are built through a layer-by-layer deposition process on a glass or silicone substrate using this system. The PAM system also permits selective placement of cells and proteins in scaffolds. It offers cellular-scale resolution, which is

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Table 13.3 Near net shaping RP processes with feeding systems Technology

Feed

Resolution (μm)

Selective laser melting Three dimensional printing Fused deposition modeling Robocasting LAM (POM, Trumpf, LENS™) Electron beam – ARCAM Pressure-assisted microsyringe 3D fiber deposition Electron beam – Sciaky Plasma transferred arc – MER TheriForm™ Ultrasonic – Solidica Pulsed gas tungsten arc welding – Liburdi Pulsed gas metal arc welding – EWI

Powder bed Powder bed Filament Organic ink Powder Powder bed Viscous solution Filament Wire Wire/powder Powder form Tape/powder Wire Wire/powder

500 200 250 100–1000

10–600 250

300

remarkably higher than the other techniques described previously. Overall, the process is quite similar to soft lithography.

13.4.8 Precision extrusion deposition (PED) The PED system uses an extruder which is equipped with a built-in heating unit to melt the feedstock material, which is a polymer. This eliminates the needs and problems associated with creating precursor filament with essential extrusion properties, as in fused deposition. Table 13.3 presents a list of near net shaping RP systems with their feeding system and process resolution.

13.5

Designing structural implants

Skeletal designs and structures in vertebrates are believed to represent ideal solutions that meet the functional and structural demands for survivability. It is thought that through the process of evolution, the skeletal system of mankind and other vertebrates developed with optimal designs and structures. However, a closer look might reveal that the structural design development through evolution may not be optimal, and does vary between individuals of the same species. Anatomical structures seen in living organisms today were developed through long processes that led to perfection (if we believe so) in functionality. However, the structural development in vertebrates is arguable. This brings a philosophical query: Is design concept an important issue for consideration in scaffold making?

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Also, does design profoundly impact both the short- and long-term survivability and functionality of scaffolds? It can be argued, with evidence, that during the last several hundred years, science and engineering have developed a coherent approach to evaluating the quality of mechanical structures through proper design and development. So, why should we not use what we already know? Engineers have used mechanical design to fabricate complex functional structures for a variety of applications. Though evolutionary biologists tend to consider the concept of design as being teleological and the science of biological design development and mechanical design fabrication as being poles apart, considering our vast engineering knowledge, it is rational to use mechanical design as a useful tool in developing tissue engineering scaffolds. Demands of current biomedical challenges and the increased life expectancy of mankind demonstrate the need for replacement of natural biological tissue specific to each patient to cure irreversible damage by pathogens, injury/accidents and/or physiological deformity. Currently, this need is fulfilled through the use of several kinds of biomaterials which provide a solution without scarcity of supply – an alternative to the ‘gold standard’ of tissue grafting. The orthopedic market in the United States alone is growing at a steady rate of 14–16%, annually. In order to restore the lost functions in patients, use of implants has become a routine procedure for the surgeon. Mimicking the articulation of a natural joint through the use of synthetic biomaterials magnifies the complexity of prosthetic devices or implant technology. Such complexities can be addressed through unified collaborative approaches integrating advanced manufacturing and engineering practices with the biomedical issues. While the structural design of a scaffold is vital and easy to perceive, the design and development of novel biomaterials addressing the biomedical needs for specific application is equally critical and should receive equal considerations. The importance of materials design in advancing scaffold technology should not be underestimated, and should possibly be a part of an overall implant development approach. While the practicality of such an approach can be cumbersome, time-consuming and expensive, an integrated innovative approach of that kind possesses potential for radically transforming implant technology to the next level. Designing of biomaterials at different length scale levels is critical for controlling and tailoring cell– biomaterial interactions that can be used beneficially in developing implants with desired characteristic properties. An excellent example of design in joint prosthetic technology is the knee joint. Total joint arthroplasty is a medical procedure that removes all of the native mechanical components of a joint and replaces them with new synthetic components assembled together as a fully functional and long-lasting coupling between two articulating bones. Various prosthetic joints currently

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exist for the major natural joints in the body. Because the knee joint receives, absorbs and transmits some of the highest and most complex mechanical forces applied to the human body, its prosthetic replacement must offer an equivalent level of robustness and ideally retain its properties throughout the lifetime of the patient. However, none of the existing knee-joint prostheses last more than 15 years. A revision surgery for knee joints has minimal success. Over a 26 year period (1969–1995), 3234 knee arthroplasties were surveyed to observe an average lifetime between 10–15 years for modern designs.63 However, to address knee-joint replacement problems with younger patients, who will at least require a second surgery during their lifetime, better design of structure, functionality and materials will be necessary. The state of the art in modern knee-replacement does not demonstrate a satisfactory product lifetime due to the frequent need of revision surgery. Mimicking the articulation of a natural joint through the use of synthetic biomaterials magnifies the complexity of prosthesis technology and must be accommodated through advanced engineering and biomedical methods. Solutions are offered through the use of oxidized bioceramics replacing metal-alloy components, enhancing the tribological properties of ultra-high molecular weight polyethylene (UHMWPE) components and incorporating biomedical solutions such as pharmaceuticals, biomimetic debris coatings or even the use of a vitamin E dopant. Another example where proper design of a bio-structure has helped change the use of technology in clinics is artificial disc replacement – an emerging surgical technique in spinal repair. Artificial disc replacement is not a new procedure, since some designs of artificial discs have been around for over 40 years. It is estimated that 80% of Americans will experience at least one episode of serious back pain during their lifetime. For many of these people, back pain will become a serious problem having lifestyle altering effects. In the United States alone it is estimated that 60 billion dollars will be spent annually on health care costs related to the treatment of disc degeneration and its related diseases. When lost wages due to the inability to work are factored in, the economic costs become even higher. The mechanical artificial disc is provided with metal-on-metal design. The most famous design is the MAVERICK artificial disc. Advantages of metalon-metal disc design may be recognized in its high tensile strength, high fatigue strength, and high corrosion resistance. However, this design may show some weakness in stiffness of the system and lack of flexibility. Due to limitations of the mechanical and structural properties of these older discs, their use has not been widespread. With re-design of the disc, artificial disc replacement now grows to be a practical alternative for spine fusion. There are many different factors designers must keep in mind as they develop an artificial disc. The device must be able to maintain appropriate intervertebral spacing, tolerate the full range of motion, and provide stabil-

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ity. Disc materials must be biocompatible, be able to form a strong interfacial bond, and perform intended functions without failure for the lifetime of the patient. Most of the current discs (e.g. PRODISC®, approved for clinical study by FDA) are made of a bioinert polymer sandwiched between metallic plates. They must come in a variety of sizes to accommodate patient height and spacing needs. Besides this, the artificial disc must be very durable. This means that the implant must last at least 50 years, since today the average age of a patient needing a lumbar disc replacement is about 35 years. Total disc replacement (Link® SB Charité III) and disc nucleus replacement are the two major methods developed. The FDA (Food and Drug Administration) has approved the CHARITÉ® Artificial Disc for use in treating pain associated with degenerative disc disease. The CHARITÉ® Artificial Disc is comprised of proven orthopedic materials; cobalt–chromium endplates and an ultra-high molecular weight polyethylene (UHMWPE) sliding core. Practical testing indicates that the unique mobile core design incorporates a floating center of rotation (FCOR), enabling independent translation and rotation. This design helps align the spine and allows movement. The hydraulic artificial discs have a gel-like core covered with a tightly woven polyethylene jacket. Before implantation, the pallet shaped core is compressed and dehydrated to minimize its size. After implantation, the woven jacket allows fluid to pass through the core in a period of 24 hours. It takes 4 to 5 days for maximum expansion of the hydraulic artificial disc. Although the mechanical goals of motion preservation using artificial discs versus motion elimination with fusion are diametrically opposed, the clinical goals of decreased pain and increased function remain intact. Artificial disc replacement has the potential of revolutionizing spinal procedures. The concept of restoring motion against elimination can prove more beneficial in maintaining the posture and structural features of a natural spine. However, this emerging technology requires extensive research to create materials designed to mimic the structure and functions of natural discs. Current practice in device design usually starts from biomechanical analysis of the stress distribution and functionality of a particular device. One of the main issues in orthopedic implant design is the fabrication of scaffolds that closely mimic the biomechanical properties of the surrounding bone. If it is a joint related device, it is important that the patient can actually move the joint along multiple directions and planes to properly restore and recover its functionality. Figure 13.4 schematically presents the important issues to be considered during skeletal device designs. CAD/CAM and RP technologies have played a significant role in orthopedic device design. Using this approach, real information from patients can be gathered using

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Materials issues • Composition • Structure/phase • Surface properties • Mechanical properties • Biological properties

373

Design issues • Geometry • Shape and contour • Functionality • Process feasibility Application • Patient range • Defect types and sizes • Longevity • Storage and sterilization

Orthopedic device

13.4 Orthopedic device – issues of importance in design, development and fabrication.

a computed tomography (CT) or magnetic resonance imaging (MRI) scan, which then can be regrouped to form a three-dimensional image using commercial software. This 3D data can be transformed to a CAD file and reworked using other available software to address the needs. The CAD file can be used to redesign or modify orthopedic devices that will be suitable to perform the patient’s needs. If necessary, the device can also be tested in a virtual world using finite element analysis (FEA) to optimize its functionality. Optimized device can then be fabricated using mass manufacturing technologies. Figure 13.5 shows a schematic of new design and development approach for creating structural implants, using advanced imaging and RP/advanced manufacturing.

13.6

Rapid prototyping in biomedical engineering – synopsis

Tissue engineering unites the science and technology of cell biology, engineering and biomaterials, along with suitable biochemical and physicochemical factors to either replace, rejuvenate or improve damaged or diseased tissue. Tissue engineering aims to take cells from a patient, expand their numbers to a desired level and seed them on a scaffold. New tissue is formed over the scaffolds within a short time through the applications of growth factors or stimuli, before being implanted in the patient to restore functions. Today, in most cases, the scaffolds are directly implanted in vivo without cell seeding. The scaffold is a three-dimensional substrate and it serves as a template for tissue regeneration. The scaffold provides a frame-

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Design and development of structural implants Stage 1 Imaging of damaged tissue • Magnetic resonance imaging (MRI) • Computer tomography (CT) Image reconstruction – good tissue

Stage 2 Modeling (CAD/CAM/ProE etc.) • Design to mimic internal tissue geometry • Design to mimic external structure (from Stage 1)

Stage 3 Design and analysis • Virtual reality (VR) design • Analysis of VR design • FEM modeling and analysis

Stage 4 Fabrication of implants • RP-based techniques for evaluation • Advanced manufacturing techniques for commercialization

13.5 Schematic of a new design and development approach for structural implants.

work and initial support for the cells to attach, proliferate and differentiate. One of the objectives of tissue engineering is to develop better scaffolds for diseased or damaged tissues that require replacement or repair through an integrated multidisciplinary collaborative approach of engineering, biology and medicine. As scaffolds are necessary for cells to grow, develop and regenerate, designing of scaffolds with suitable architecture, surface properties and strength to serve their intended function and purpose has been one of the core challenging areas for biomedical engineers. Tissue engineering scaf-

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Compact bone Spongy bone

13.6 Schematic of compact and spongy bone showing interconnected porous structure.

folds are developed to perform the following functions: (a) to serve as an adhesion substrate for the cell, facilitating the localization and delivery of cells when they are implanted; (b) to provide temporary mechanical support to the newly grown tissue by defining and maintaining a 3D structure; and (c) to guide the development of new tissues with the appropriate function.64 However, the functions and the needs of a scaffold also depend on its application. Scaffolds for structural tissue engineering or bone tissue engineering should ideally possess the following desirable characteristics: (a) be osteogenic and resorbable with a resorption rate matching the repair process; (b) have an interconnected porous structure with pores in the range of 150–400 μm to promote bone in-growth; (c) have suitable surface chemistry for cell attachment, proliferation and differentiation; and (d) have matching mechanical properties to that of the tissue being replaced. Figure 13.6 shows a schematic of cancellous (spongy) and cortical (compact) bone having interconnected porous structure. Looking at the natural composition of bone, it is believed that one of the promising approaches is to use 3D porous ceramic and/or polymer–ceramic composite scaffolds that have structural and compositional resemblance with bone. Although tissue grafting techniques (autograft, isograft, allograft and xenograft) have been known for a very long time, there are serious disadvantages associated with these techniques which include donor site morbidity, mechanical strength limitations, limited availability of tissue for grafting,

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need for multiple surgeries, transmission of diseases and rejection by the patient’s immune system triggered by an immunological response. For these reasons, there is a growing demand for design and development of artificial scaffolds for tissue engineering applications. In addition to materials, processing and manufacturing issues, designing of scaffolds for tissue repair to mimic structural intricacies and functionality, pose serious challenges for traditional manufacturing approaches. However, several RP processes have shown significant potential for developing patient-specific scaffolds with different structural properties using several kinds of materials. Applications of RP in tissue engineering offer the pledge of growing regenerative tissue and functional organs, which in the course of time become biologically indistinguishable. Although the first RP technique, available in the late 1980s, were primarily used to produce models and prototypes only, RP technologies are today used to create functional structures for several types of engineering and biomedical applications. This approach, when integrated with medical science and clinical practices will definitely improve the health care system and quality of human life. One of the main goals of scaffold technology in biomedical engineering is to closely mimic the biomechanical properties of the tissues to be replaced, regenerated or rejuvenated. Conventional manufacturing techniques are not satisfactory for fabricating scaffolds with controlled internal architecture to modulate mechanical properties. Fortunately, during the last decade, RP technologies have emerged as a convenient low volume or one-of-akind part manufacturing technique in which parts can be directly fabricated from their computer-aided design (CAD) data description without using any part-specific tooling or dies. Advances in both computational topology design (CTD) and SFF have made it possible to create scaffolds with controlled architecture, both macroscopically and microscopically. During recent years, RP techniques have attracted significant interest in biomedical engineering applications as powerful tools for scaffold fabrication. In the medical field, RP has been used to create tools, implants including bioartificial liver65 and prosthetic devices, as well as bio-models.47,49,50,62,66–68 Other processing techniques such as SLS69–72 and stereolithography73–80 have already been used to fabricate controlled porosity ceramic implants for orthopedic applications. Scaffolds fabricated through RP techniques are built through a layer-by-layer manufacturing process during which material/s is/are deposited on a stage/bed, either in a molten phase (commonly referred to as FDM),45,51,68,81,82 or in droplets together with a binding agent (referred to as 3D Printing).49,83,84 These techniques can create 100% interconnected 3D porous scaffolds with controlled pore size, pore shape and geometry. RP techniques provide great flexibility and versatility in creating scaffolds that can mimic both the internal and external structures of the hard tissue of interest.

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RP methods such as 3D plotting62 and 3D fiber deposition84,85 have been developed recently and are being used in tissue engineering. RP or SFF techniques can also fabricate biphasic scaffolds that incorporate multiple geometries into a single scaffold,86 allowing growth of more tissues into a single scaffold structure. It is also possible to integrate ‘cell seeding’ during the scaffold fabrication process, thereby minimizing the problems with poor cell infiltration and in-growth into the scaffold, which is a common problem with most of the conventional techniques.87 The advantages of RP or SFF technology currently in use in tissue repair have been reviewed by several authors.50,51,62,83,88 Some recent researchers have displayed a new outlook on the use of RP in biomedical engineering. They believe in developing organs directly, using the concept of layered manufacturing techniques – RP processes. These new technologies include laser printing of cells,89 organ printing,90,91 photopatterning of hydrogel92 and microfluidics technology.92 Emergence of these kinds of new approaches in scaffolds, as well as 3D organ fabrication through layer-by-layer manufacturing techniques, underline the truth that the use of advanced manufacturing engineering technology in tissue engineering is growing, which is evident from the growth in scientific literature on that topic during recent years. Table 13.4 presents a list of recent review articles discussing applications of RP and SFF in tissue engineering.

13.7

Rapid prototyping in mimicking structural intricacies of biological materials

The ability to control architectural features in tissue engineering scaffolds is crucial for tissue regeneration and long-term mechanical, physiological and biological performance of the scaffolds. Significant research efforts have been made to mimic internal architecture of tissues, particularly of bone, using simple as well as advanced techniques. Although a good amount of success has been made in developing porous metallic, ceramic and composite scaffolds for bone tissue repair, the current technology has to go a long way before mankind can mimic natural tissue in entirety – functionally, compositionally and in architecture. It is demonstrated that pores of less than 10 μm inhibit cellular ingrowth, while pores between 15 and 50 μm help fibrovascular colonization, pores between 50 and 150 μm help osteoid growth, and pores higher than 150 μm facilitate internal mineralized bone formation. RP techniques provide great flexibility and versatility in creating scaffolds that can mimic both the internal and external architecture of the tissue of interest. RP techniques including 3D plotting,62 FDM45,47,48,68,99 and 3D fiber deposition,84,85 developed during recent years, have successfully created

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Table 13.4 Recent review articles on applications of RP or SFF in tissue engineering Year

Title

Authors and reference number

2008

Engineering functionally graded tissue engineering scaffolds

2008

Scaffold design and fabrication, in Tissue Engineering

2007

Design and development of three-dimensional scaffolds for tissue engineering Rapid prototyping as a tool for manufacturing bioartificial livers Rapid manufacturing of metal components by laser forming Porous scaffold design for tissue engineering Rapid prototyping in tissue engineering: challenges and potential Three-dimensional tissue fabrication Solid freeform fabrication of three-dimensional scaffolds for engineering replacement tissues and organs Making tissue engineering scaffolds work. Review on the application of solid freeform fabrication technology to the production of tissue engineering scaffolds The design of scaffolds for use in tissue engineering. Part II. Rapid prototyping techniques Scaffolds in tissue engineering bone and cartilage

K.F. Leong, C.K. Chua, N. Sudarmadji, and W.Y. Yeong93 D. Hutmacher, T. Woodfield, P. Dalton, J. Lewis, B. Clemens van, T. Peter, L. Anders, H. Jeffrey, F.W. David, C. Ranieri, D.d.B. Joost, and S. Jérôme51 C. Liu, Z. Xia, and J.T. Czernuszka94

2007 2006 2005 2004

2004 2003

2003

2002

2000

X. Wang, Y. Yan, and R. Zhang65 E.C. Santos, M. Shiomi, K. Osakada, and T. Laoui58 S.J. Hollister95 W.Y. Yeong, C.K. Chua, K.F. Leong, and M. Chandrasekaran50 V. Liu Tsang and S.N. Bhatia96 K.F. Leong, C.M. Cheah, and C.K. Chua67

E. Sachlos and J.T. Czernuszka43

S. Yang, K.F. Leong, Z. Du, and C.K. Chua97 D.W. Hutmacher98

scaffolds for tissue repair. The 3D fiber deposition technique has shown great potential in tissue engineering through its precision 3D scaffold making, characterized by 100% interconnected pores with different shapes and sizes.84 Next, the author describes some of the RP techniques explored by researchers to mimic structural intricacies of biological materials (tissues and organs).

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13.7.1 Fused deposition modeling (FDM) in scaffold technology FDM technology provides unique advantages in mimicking structural intricacies of biological materials in scaffold technology. The FDM process allows one to control the shape and size, as well as the internal architecture, of the porous scaffolds; which can be easily designed using CAD/CAM, ProE, Insight® or similar software. The fundamental FDM processing parameters that govern shape, size and volume of pores in the final part/ structure are road or raster width, road gap, slice thickness and the angle between two successive layers of roads. Figure 13.7 shows a wax ‘3D honeycomb’ porous prototype made with basic FDM processing parameters. FDM, being a layer-by-layer manufacturing technique, can be used to manufacture prototypes in which each layer has different road-width, road-gap and angle between successive layers. These processing parameters can be easily altered to create structural tissue engineering scaffolds with desired pore size, geometry, interconnectivity and biomechanical performance. This novel concept of designing and developing gradient-controlled porosity structures with complex internal architectures can be used to fabricate bone implants similar to human bone, especially for cancellous bone graft application. Strength, biocompatibility, biodegradability, and bioresorbability of these implants will mainly depend on the choice of materials.

Inches

13.7 ‘3D honeycomb’ porous prototype (mold design) made with basic FDM processing parameters.

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Pore size and pore volume in these scaffolds can be tailored as per requirements, which also influences mechanical strength and rate of biodegradation to some extent. Design and selection of proper FDM processing parameters are the key tools in achieving scaffolds with complex internal geometry, while selection of appropriate materials in scaffold fabrication will decide the biological and interfacial properties. The FDM process has also been used indirectly to create porous ceramic scaffolds for hard tissue engineering applications. The indirect fused deposition process fabricates porous polymeric molds using an FDM machine. The molds are then infiltrated with the ceramic slurry of interest, followed by binder-removal and sintering to produce porous ceramic scaffolds. The final scaffold is the negative of the mold designed. Therefore, by altering the FDM processing parameters during mold design (e.g. road width, road gap and angle between successive layers), the pore size, and pore volume in the final scaffold can be tailored. Kalita et al. (2002) developed porous calcium aluminate scaffolds with varying volume fractions of porosity (29 vol% and 44 vol%) and pore size of 300 μm, by tailoring the FDM processing parameters. FDM processing parameters used to develop wax molds to fabricate porous ceramic structures with 29 vol% porosity and 305 μm pore size were 0.41 mm road width, 1.01 mm road gap, and 0.25 mm slice thickness. For 44 vol% porosity structures, the road gap was changed to 0.51 mm keeping the other parameters constant. These controlled porosity wax molds were infiltrated with ceramic slurry and then sintered to develop cylindrical porous calcium aluminate scaffolds.47 Through the direct FDM route, Zein et al. (2002) fabricated porous polycaprolactone (PCL) scaffolds with a honeycomb structure having a channel size of 160–770 mm.45 Kalita et al. (2003) successfully designed, developed and fabricated 3D interconnected controlled porosity polypropylene– tricalcium phosphate (PP–TCP) composite scaffolds via the direct FDM process.68 PP–TCP composite filaments were processed using a single-screw extruder followed by fused deposition of these filaments to fabricate porous structures.68 Samples had a pore size of 160 μm and pore volumes of 36%, 40%, and 52%, and a highest compressive strength of 12.7 MPa was reported, which is comparable to the tensile strength of natural cancellous bone (7.4 MPa).68 The scaffolds were non-toxic and osteo-precursor cells were attached and anchored well on the surface of these composite scaffolds.68 Figure 13.8 shows three-dimensionally, interconnected, controlled porosity PP–TCP composite scaffolds with different internal architectures fabricated using the FDM process.68 Endres et al. (2003) investigate the in vitro osteogenic potential of human mesenchymal progenitor cells in novel matrix architectures built by FDM techniques using polycaprolactone (PCL) and polycaprolactonehydroxyapatite (PCL-HA) materials.100 This scaffold design concept was

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Internal architecture

3D honeycomb porous PP–TCP composite scaffolds

Internal architecture

13.8 Three-dimensionally interconnected controlled porosity PP–TCP composite scaffolds with different internal architectures fabricated using the fused deposition modeling process. (Reprinted from Materials Science and Engineering C, 23, 611–620, S.J. Kalita, S. Bose, H.L. Hosick, and A. Bandyopadhyay, ‘Development of controlled porosity polymer–ceramic composite scaffolds via fused deposition modeling’, Copyright (2003), with permission from Elsevier.68)

characterized by novel matrix architecture, good mechanical properties and slow degradation, and a biomimetic milieu for cell delivery and proliferation.100

13.7.2 Stereolithography in scaffold technology Melchels et al. (2009) designed and prepared porous polylactide constructs by stereolithography, for the first time, without the use of reactive diluents. Star-shaped poly(D,L-lactide) oligomers with 2, 3 and 6 arms were synthesized, end functionalized with methacryloyl chloride and photo-crosslinked in the presence of ethyl lactate as a non-reactive diluent.78 Films and porous

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13.9 Images of PDLLA network scaffolds with a gyroid architecture, built by stereolithography. (a) Photograph, (b) mCT visualisation and (c) SEM image. In panel (d) a light microscopy image is shown for a scaffold seeded with mouse pre-osteoblasts after 1 day of culturing. Scale bars represent 500 μm. (Reprinted from Biomaterials, 30, 3801–3809, F.P.W. Melchels, J. Feijen, and D.W. Grijpma, ‘A poly(d,l-lactide) resin for the preparation of tissue engineering scaffolds by stereolithography’, Copyright (2009), with permission from Elsevier.78)

scaffolds with gyroid architecture were prepared by stereolithography, using a liquid resin based on a two-armed PDLLA macromer and ethyl lactate.78 Mouse pre-osteoblasts readily adhered and proliferated well on these networks. Open pore architecture of these hydrophobic scaffolds, built as per pre-designed (porous) architectures as shown in Fig. 13.9, helped in cell seeding.78

13.7.3 Three-dimensional printing in mimicking structural intricacies Chen et al. (2006) of the University of Michigan developed computerdesigned nano-fibrous scaffolds with complex geometries on the macroand micro-scales.101 They used reverse SFF and thermal phase separation of poly(L-lactic acid) (PLLA) solutions to create three-dimensional nanofibrous scaffolds.101 Figure 13.10 shows their three-dimensional PLLA scaffolds, created from reverse solid-freeform fabrication.101 This approach allows for the fabrication of nano-fibrous matrices while having precise

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(a)

(b)

(c)

13.10 Three-dimensional PLLA scaffolds created from reverse solidfreeform fabrication. (a) Negative mold design used for scaffold casting. Solid struts in the mold eventually become the open pores in the final PLLA scaffold. (b) SEM micrograph of the overview of a nano-fibrous scaffold showing micro-pore structure in the struts. (c) SEM micrograph of the fibrous morphology of the nano-fibrous scaffold pore walls. Scale bars of (b) and (c) are 500 μm and 2 μm, respectively. (Reprinted from Biomaterials, 27, 3973–3979, V.J. Chen, L.A. Smith, and P.X. Ma, ‘Bone regeneration on computer-designed nano-fibrous scaffolds’, Copyright (2006), with permission from Elsevier.101)

control of internal pore size and structure, as well as external scaffold shape including architectures generated from computed-tomography scans and histological sections.101 The ability to control these features is critical for the success of neo-tissue regeneration. They used Rhinoceros software, Modelworks software (Solidscape) and a 3D printing system Modelmaker II (Solidscape) to create these scaffolds for bone tissue engineering.101 Threedimensionally interconnected porous scaffolds with nano-scale morphology are suggested to affect cell behavior in several tissues.102,103

13.7.4 Selective laser sintering Williams et al. developed porous polycaprolactone (PCL) polymeric scaffolds that were computationally designed and then fabricated via selective laser sintering (SLS), a RP technique.104 PCL is a bioresorbable polymer with potential applications for bone and cartilage repair. PCL scaffolds

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fabricated via SLS show great potential for replacement of skeletal tissues.104 These scaffolds possess mechanical properties within the lower range of cancellous or trabecular bone; the mechanical properties can be computationally analyzed, bypassing the need for experimental testing.104 Vail et al. (1999) developed and characterized porous calcium phosphate ceramic structures using SLS techniques and showed that these scaffolds were biocompatible.105

13.7.5 Three-dimensional fiber deposition Woodfield et al. (2004) developed a fiber deposition technique for producing three-dimensional poly(ethylene glycol)terephthalate–poly(butylene terephthalate) (PET/PBT) block copolymer scaffolds with a 100% interconnecting pore network for the engineering of articular cartilage.85 The 3D fiber deposition technique allowed them to ‘design-in’ desired scaffold characteristics, such as porosity, pore size and mechanical properties, using computer-controlled tooling processes.85 Li et al. (2006) used three-dimensional fiber deposition to produce novel 3D porous Ti-6Al-4V scaffolds with fully interconnected porous networks and highly controllable porosity and pore size.106 They showed that this technique could deposit Ti-6Al-4V slurry at room temperature to produce a scaffold consisting of layers of directionally aligned Ti-6Al-4V fiber, based on a 3D computer-controlled strategy.106 The novel scaffold fabrication processes of 3D fiber deposition exhibit great potential in making reproducible 3D scaffolds with precision, to control scaffold structure to modulate the mechanical and biological properties.

13.7.6 Robotic dispensing Ang et al. (2002) designed a new robotic desktop RP system to fabricate scaffolds for tissue engineering applications using chitosan and chitosanHA dissolved in acetic acid. They demonstrated that a RP robotic dispensing (RPRD) system has potential in fabricating three-dimensional scaffolds with regular and reproducible macropore architecture.107 The produced scaffolds were biocompatible.107

13.7.7 Electron beam melting for 3D porous scaffolds Harrysson and Cormier (2005) described the EBM fabrication of customized orthopedic implants from commercially available powders of H13 steel and Ti-6Al-4V.108 In a recent report Murr et al. (2009) described an exploratory characterization and comparison of EBM Ti-6Al-4V components (from nominal 30 μm diameter powder) with wrought products. EB

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manufactured product with an average hardness of 3.6–3.9 GPa, compared to 4.0 GPa for wrought products.109 They also discussed biomedical applications of EBM prototypes in direct prosthesis or implant manufacturing from CT or MRI data, especially from the viewpoint of tailoring physical properties through EB control to achieve customized and optimized implant and prosthetic products.109 In another recent publication, Li et al. (2009) utilized EBM technique to fabricate porous Ti-6Al-4V parts with fully interconnected, controlled internal pore architecture.110 They demonstrated the potential of EBM process for fabricating Ti-6Al-4V parts with internal pore architecture meeting the requirements of orthopedic implants.110

13.7.8 Other customized rapid prototyping processes for mimicking structural intricacies Porous PCL and PLLA scaffolds with line width of 20 mm were developed by Vozzi et al. using PAM technique.111 Porous PLLA scaffolds with controllable porous architectures, having sizes in the range of 200 to 500 μm, were produced and characterized by a new technique called precision extrusion manufacturing (PEM).112 Ryan et al. (2008) reported a multi-stage RP technique that was developed to successfully produce porous titanium scaffolds with fully interconnected pore networks and reproducible porosity and pore size.113 The scaffolds’ porous characteristics were governed by a sacrificial wax template, fabricated using a commercial 3D printer, while a powder metallurgy route was employed to generate the scaffolds.113 Lopez-Heredia et al. (2008) reported the development of porous titanium scaffolds with a regular array of interconnected pores of 1000 μm diameter and a porosity of 50%, using a RP technique followed by a calcium phosphate coating using an electrodeposition method.114 Their work demonstrated that porous Ti scaffolds with high strength and defined geometry can be evenly coated with calcium phosphate layers and cultured mesenchymal stem cells for bone tissue engineering.114 Control of the internal architecture of scaffolds is essential to create complex organs. Xu et al. (2008) designed and fabricated complex 3D objects with an intrinsic network of interconnected channels, using PCL and PEG, to create vascular systems that mimic human livers by using a lowtemperature deposition system.115 The design and fabrication strategies used directly from CAD models present a promising route for making complex organ vascular systems.115 Some of the recent progress involving the use of RP/SFF technology in mimicking structural intricacies of biomedical materials is summarized in Table 13.5.

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Table 13.5 Progress of RP or SFF in tissue engineering Year

RP/SFF process

Research and development

Ref.

2009

EBM

109

2009

EBM

2008

Low temperature deposition

2006

3D fiber deposition

2005

EBM

2005

Stereolithography

2004

3D fiber deposition

2003

Direct FDM

2003

Direct FDM

2003

PAM

2003

Inkjet printing

2002

Direct FDM

2002

Indirect FDM

2002

3D plotter

2002

Robotic dispensing

2002

3D printing

2001

PEM

1999

SLS

1999

Indirect FDM

Ti-6Al-4V biomedical prototypes compared to wrought Ti-6Al-4V Interconnected porous Ti-6Al-4V parts with controlled internal pore architecture Complex 3D objects using PCL and PEG, to create vascular systems that mimic human livers Ti-6Al-4V scaffolds with fully interconnected porous networks and highly controllable porosity and pore size Customized orthopedic implants from H13 steel and Ti-6Al-4V commercial powders Biomimetic porous scaffolds of calcium phosphate with a strut size of 450 μm 3D PEGT/PBT block co-polymer scaffolds with 100% interconnected pores for articular cartilage 3D porous PP-TCP composite biocompatible scaffolds Cell seeded PCL and PCL-HA 3D porous scaffolds Porous PCL and PLLA scaffolds with line width of 20 mm Inkjet printing system for cells or biomaterials. 3D organ printing which accelerates and optimizes tissue and organ assembly Honeycomb PCL scaffolds with channel size of 160–770 mm Porous calcium aluminate scaffolds with varying pore volume Thermoreversible hydrogels with 200–300 μm interconnected pores 3D chitosan–hydroxyapatite scaffolds Porous scaffolds were created by 3DP process by developing a unique blend of starch-based polymer powders (cornstarch, dextran and gelatin) Porous PLLA scaffolds with a new RP technique Porous calcium phosphate ceramic structures 3D porous controlled porosity ceramic scaffolds

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110 115 106

108 116 85

68 100 111 90, 117, 118

45 47 62 107 49

112 105 119

Rapid prototyping in biomedical engineering

13.8

387

Patient-specific customized scaffolds via rapid prototyping

RP or SFF techniques possess the capability of designing and fabricating anatomically shaped scaffolds for the damaged or diseased tissue of interest with varying internal architectures, while perfectly mimicking the external features. This helps to control pore size, pore volume, permeability, pore connectivity, strength and stiffness, with greater precision. Control over pore architecture and interconnectivity enhances cell attachment, cell infiltration and mass transport, while the ability to fabricate scaffolds with the patients’ anatomical shape for the tissue of interest provides the capability of fabricating patient-specific customized scaffolds for tissue repair. Although there are numerous challenges associated with making the dream of ‘patient-specific customized scaffolds’ a reality, advances in RP technology and tissue engineering show promise. In future, scaffolds will be fabricated on site (in clinics, or in close proximity) within a very short time span, under the guidance of cross-disciplinary experts. Ability to remotely control a RP machine over the worldwide web is very exciting. Therefore, with proper integration and coordination, it is feasible to unite engineers, doctors, surgeons and biomaterial scientists, in remote locations, to create patientspecific customized scaffolds, to bring our dream to reality. Chen et al. (2006) of the University of Michigan showed the feasibility and versatility of a reverse SFF to mimic patient-specific human anatomical parts using CT scans or histological sections.101 They created porous anatomical parts of PLLA using 3D printing, as shown in Fig. 13.11. Williams et al. demonstrated the feasibility of manufacturing PCL scaffolds via the SLS technique, to fit complex anatomic locations.104 They fabricated a pig mandibular condyle scaffold using this approach.104 Figure 13.12 shows the development of patient-specific resorbable ceramic mandible scaffolds using the indirect FDM technique. A comparison of different RP techniques in manufacturing of bioartificial livers is published by Wang et al. in Trends in Biotechnology.65

13.9

Conclusion

Demand for an integrated approach of biology and engineering is growing, which is evident from the growth of interdisciplinary technologies. RP technologies hold a promising future with great potential in scaffold, implant and future artificial organ technology. RP, combined with advanced imaging technology and analysis software, not only can create scaffolds with controlled design and architecture, but also can create patient-specific customized scaffolds and implants, and analyze them virtually for properties. Indirect or customized RP methods, coupled with conventional

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13.11 Scaffolds created from 3D reconstructions of CT scans or histological sections. (a) Human ear reconstruction from histological sections and (b) the resulting nano-fibrous scaffold (scale bar, 10 mm); (c) human mandible reconstruction from CT scans, segment shows the reversed image of the bone fragment to be engineered; (d) resulting nano-fibrous scaffold of the mandible segment (scale bar, 10 mm); (e) SEM micrographs of the interconnected spherical pores within the mandible segment (scale bar, 500 μm); (f) the nano-fibrous pore morphology of a spherical pore (scale bar, 5 μm). (Reprinted from Biomaterials, 27, 3973–3979, V.J. Chen, L.A. Smith, and P.X. Ma, ‘Bone regeneration on computer-designed nano-fibrous scaffolds’, (2006) Copyright (2006), with permission from Elsevier.101) © Woodhead Publishing Limited, 2010

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13.12 Development of patient-specific resorbable ceramic mandible scaffolds. (a) Human jaw; (b) STL file of the human jaw created from CT scan data; (c) object file of the jaw with support; (d) cross-sectional view of mandible; (e and f) porous mandible mold created in FDM with ABS filament infiltrated with tricalcium phosphate ceramic slurry.

pore-forming techniques expand the base and applications of RP in mimicking structural intricacies of structural materials. Future advancement of RP in tissue engineering entails the need for new materials, better software and imaging techniques, innovative and collaborative efforts from cell physiology and materials scientists, ceramicists, polymer chemists, mechanical engineers, molecular biologists, surgeons and other medical professionals.

13.10 List of abbreviations 3DFD 3DP CAD CAM CTD DLF DMD DMLS EBM ECM FDM FPD

three-dimensional fiber deposition three-dimensional printing computer aided design computer aided manufacturing computational topology design direct light fabrication direct metal deposition direct metal laser sintering electron beam melting extracellular matrix fused deposition modeling functional prototype development

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HA LAM LENSTM MEMS NEMS PAM PBT PCL PDLLA PED PEM PET PGA PHB PLA PLLA PMMA PP RM RP SEM SFF SLM SLS SMA TCP THR

hydroxyapatite laser additive manufacturing laser-engineered net shaping micro-electromechanical system nano-electromechanical system pressure-assisted micro-syringe poly(butylene terephthalate) polycaprolactone poly(dl-lactic acid) pressure extrusion deposition precise extrusion manufacturing poly(ethylene glycol) terephthalate poly(glycolic acid) or polyglycolide poly(3-hydroxy butyrate) poly(lactic acid) or polylactide poly(L-lactic acid) poly(methyl methacrylate) polypropylene rapid manufacturing rapid prototyping scanning electron microscopy solid freeform fabrication selective laser melting selective laser sintering shape memory alloy tricalcium phosphate total hip replacement

13.11 References 1 ‘The Bone and Joint Decade 2000–2010 for prevention and treatment of musculoskeletal disorders’, Acta Orthop Scand; suppl. 281, 69 (1998). 2 k. annan, The Secretary General Message to Launch The Bone and Joint Decade 2000–2010 for the Prevention and Treatment of Musculoskeletal Disorders [letter]. New York, UN Headquarters, Nov 30, 1999. 3 r.e. coleman, ‘Metastatic bone disease: Clinical features, pathophysiology, and treatment strategies’, Cancer Treat Rev, 27, 165–76 (2001). 4 t. yoneda, ‘Cellular and molecular mechanisms of breast and prostate cancer metastasis in bone’, Eur J Cancer, 34, 240–245 (1998). 5 f.w. orr, s. sanchez, o.h. kostenuik, and p.g. singh, ‘Tumor–bone interactions in skeletal metastasis’, Clin Orthop Related Res, 312, 19–33 (1995). 6 p.i. branemark, ‘Osseointegration and its experimental background’, J Prosthet Dent, 50, 399–410 (1983). 7 r. shalak and c.f. fox, Preface, in Tissue Engineering, R. Shalak and C.F. Fox, Editors. 1998, Alan R. Liss: New York. p. 26–29.

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8 r. langer and j.p. vacanti, ‘Tissue engineering’, Science, 260, 920–926 (1993). 9 k. whang, c.k. thomas, g. nuber, and k.e. healy, ‘A novel method to fabricate bioabsorbable scaffolds’, Polymer, 36, 837 (1995). 10 l.g. cima, j.p. vacanti, c. vacanti, d. inger, d. mooney, and r. langer, ‘Tissue engineering by cell transplantation using degradable polymer substrates’, J Biomech Eng T ASME, 113, 143–151 (1991). 11 a.g. mikos, y. bao, l.g. cima, d.e. ingber, j.p. vacanti, and r. langer, ‘Preparation of poly (glycolic acid) bonded fibres structures for cell attachment and transplantation’, J Biomed Mater Res, 27, 183–189 (1993). 12 l.e. freed, g. vunjak-novakovic, r.j. biron, d.b. eagles, d.c. lesnoy, s.k. barlow, and r. langer, ‘Biodegradable polymer scaffolds for tissue engineering’, Biotechnology (N Y), 12, 689–93 (1994). 13 y.y. hsu, j.d. gresser, d.j. trantolo, c.m. lyons, p.r. gangadharam, and d.l. wise, ‘Effect of polymer foam morphology and density on kinetics of in vitro controlled release of isoniazid from compressed foam matrices’, J Biomed Mater Res, 35, 107–16 (1997). 14 i.v. yannas, j.f. burke, p.l. gordon, c. huang, and r.h. rubenstein, ‘Design of an artificial skin. Part II. Control of chemical composition’, Biomaterials, 14, 107–131 (1980). 15 c.j. doillon, c.f. whyne, s. brandwein, and f.h. silver, ‘Collagen-based wound dressings: Control of the pore structure and morphology’, J Biomed Mater Res, 20, 1219–1228 (1986). 16 h. schoof, j. apel, i. heschel, and g. rau, ‘Control of pore structure and size in freeze-dried collagen sponges’, J Biomed Mater Res-A, 58, 352–357 (2001). 17 s.v. madihally and h.w.t. matthew, ‘Porous chitosan scaffolds for tissue engineering’, Biomaterials, 20, 1133–1142 (1999). 18 r. glicklis, l. shapiro, r. agbaria, j.c. merchuk, and s. cohen, ‘Hepatocyte behavior within three-dimensional porous alginate scaffolds’, Biotechnol Bioeng, 67, 344–353 (2000). 19 d.j. mooney, d.f. baldwin, n.p. suh, j.p. vacanti, and r. langer, ‘Novel approach to fabricate porous sponges of poly(lactic-co-glycolic acid) without the use of organic solvents’, Biomaterials, 17, 1417–1422 (1996). 20 r.c. thompson, m.j. yaszemski, j.m. powers, and a.g. mikos, ‘Fabrication of biodegradable polymer scaffolds to engineering trabecular bone’, J Biomater Sci-Polym E, 7, 23–38 (1995). 21 h. lo, m.s. ponticiello, and k.w. leong, ‘Fabrication of controlled release biodegradable foams by phase separation’, Tissue Eng, 1, 15–28 (1995). 22 j. guan, k.l. fujimoto, m.s. sacks, and w.r. wagner, ‘Preparation and characterization of highly porous, biodegradable polyurethane scaffolds for soft tissue applications’, Biomaterials, 26, 3961–71 (2005). 23 j. guan, j.j. stankus, and w.r. wagner, ‘Biodegradable elastomeric scaffolds with basic fibroblast growth factor release’, J Control Release, 120, 70–8 (2007). 24 m. reuber, l.s. yu, and w.j. kolff, ‘Effect of processing temperature on the properties of polyurethane and comparison of vacuum forming and solution casting to make artificial hearts’, Artif Organs, 11, 323 (1987). 25 h. okuno, k. renzo, and t. uragami, ‘Influence of casting solution additive, degree of polymerization, and polymer concentration on poly(vinyl chloride) membrane properties and performance’, J Membrane Sci, 83, 199–209 (1993).

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78 f.p.w. melchels, j. feijen, and d.w. grijpma, ‘A poly(d,l-lactide) resin for the preparation of tissue engineering scaffolds by stereolithography’, Biomaterials, 30, 3801–3809 (2009). 79 f.p.w. melchels, d.w. grijpma, and j. feijen, ‘Properties of porous structures prepared by stereolithography using a polylactide resin’, J Control Rel, 132, e71–e73 (2008). 80 k. arcaute, b. mann, and r. wicker, ‘Stereolithography of spatially-controlled multi-material bioactive poly(ethylene glycol) scaffolds’, Acta Biomater, 6, 1047–1054 (2010). 81 h. chim, d.w. hutmacher, a.m. chou, a.l. oliveira, r.l. reis, t.c. lim, and j.t. schantz, ‘A comparative analysis of scaffold material modifications for loadbearing applications in bone tissue engineering’, Int J Oral Maxillofac Surg, 35, 928–934 (2006). 82 b.c. tellis, j.a. szivek, c.l. bliss, d.s. margolis, r.k. vaidyanathan, and p. calvert, ‘Trabecular scaffolds created using micro CT guided fused deposition modeling’, Mater Sci Engng: C, 28, 171–178 (2008). 83 e. sachlos and j.t. czernuszka, ‘Making tissue engineering scaffolds work. Review on the application of solid freeform fabrication technology to the production of tissue engineering scaffolds’, Cells Mater, 5, 29–40 (2003). 84 l. moroni, j.r. de wijn, and c.a. van blitterswijk, ‘3D fiber-deposited scaffolds for tissue engineering: Influence of pores, geometry and architecture on dynamic mechanical properties’, Biomaterials, 27, 974–985 (2006). 85 t.b.f. woodfield, j. malda, j. de wijn, f. péters, j. riesle, and c.a. van blitterswijk, ‘Design of porous scaffolds for cartilage tissue engineering using a three-dimensional fiber-deposition technique’, Biomaterials, 25, 4149–4161 (2004). 86 j.m. taboas, r.d. maddox, p.h. krebsbach, and s.j. hollister, ‘Indirect solid free form fabrication of local and global porous, biomimetic and composite 3D polymer–ceramic scaffolds’, Biomaterials, 24, 181–194 (2003). 87 i. salusbury, ‘Bone in contention’, Mater World, 13, 25–27 (2005). 88 w. sun and p. lal, ‘Recent development on computer aided tissue engineering – a review’, Comput Meth Programs Biomed, 67, 85–103 (2002). 89 b.r. ringeisen, h. kim, j.a. barron, d.b. krizman, d.b. chrisey, s. jackman, r.y. auyeung, and b.j. spargo, ‘Laser printing of pluripotent embryonal carcinoma cells’, Tissue Engng, 10, 483–91 (2004). 90 v. mironov, v. kasyanov, c. drake, and r.r. markwald, ‘Organ printing: Promises and challenges’, Regen Med, 3, 93–103 (2008). 91 v. mironov, r.p. visconti, v. kasyanov, g. forgacs, c.j. drake, and r.r. markwald, ‘Organ printing: Tissue spheroids as building blocks’, Biomaterials, 30, 2164–74 (2009). 92 s.h. lee, j.j. moon, and j.l. west, ‘Three-dimensional micropatterning of bioactive hydrogels via two-photon laser scanning photolithography for guided 3D cell migration’, Biomaterials, 29, 2962–8 (2008). 93 k.f. leong, c.k. chua, n. sudarmadji, and w.y. yeong, ‘Engineering functionally graded tissue engineering scaffolds’, J Mech Behav Biomed Mater, 1, 140–152 (2008). 94 c. liu, z. xia, and j.t. czernuszka, ‘Design and Development of Threedimensional Scaffolds for Tissue Engineering’, Chem Engng Res Des, 85, 1051– 1064 (2007).

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110 x. li, c. wang, w. zhang, and y. li, ‘Fabrication and characterization of porous Ti6Al4V parts for biomedical applications using electron beam melting process’, Mater Lett, 63, 403–405 (2009). 111 g. vozzi, c. flaim, a. ahluwalia, and s. bhatia, ‘Fabrication of PLGA scaffolds using soft lithography and microsyringe deposition’, Biomaterials, 24, 2533–40 (2003). 112 z. xiong, y. yan, r. zhang, and l. sun, ‘Fabrication of porous poly(-lactic acid) scaffolds for bone tissue engineering via precise extrusion’, Scripta Mater, 45, 773–779 (2001). 113 g.e. ryan, a.s. pandit, and d.p. apatsidis, ‘Porous titanium scaffolds fabricated using a rapid prototyping and powder metallurgy technique’, Biomaterials, 29, 3625–3635 (2008). 114 m.a. lopez-heredia, j. sohier, c. gaillard, s. quillard, m. dorget, and p. layrolle, ‘Rapid prototyped porous titanium coated with calcium phosphate as a scaffold for bone tissue engineering’, Biomaterials, 29, 2608–2615 (2008). 115 w. xu, x. wang, y. yan, and r. zhang, ‘Rapid Prototyping of Polyurethane for the Creation of Vascular Systems’, J Bioactive Compatible Polym, 23, 2, 103–114 (2008). 116 a. woesz, m. rumpler, j. stampfl, f. varga, n. fratzl-zelman, p. roschger, k. klaushofer, and p. fratzl, ‘Towards bone replacement materials from calcium phosphates via rapid prototyping and ceramic gelcasting’, Mater Sci Engng: C, 25, 181–186 (2005). 117 w.c. wilson, jr. and t. boland, ‘Cell and organ printing. 1: Protein and cell printers’, Anat Rec A Discov Mol Cell Evol Biol, 272, 491–6 (2003). 118 v. mironov, t. boland, t. trusk, g. forgacs, and r.r. markwald, ‘Organ printing: computer-aided jet-based 3D tissue engineering’, Trends Biotechnol, 21, 157–61 (2003). 119 s. bose, s. suguira, and a. bandyopadhyay, ‘Processing of controlled porosity ceramic structures via fused deposition’, Scripta Mater, 41, 1009–1014 (1999).

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Index

A2780cis cell line, 211 acoustic emissions, 331 acrylic acid, 76 active scaffold, 97–101 active targeting, 247 adenosine diphosphate, 304 adipose derived stromal cells, 317 adipose tissue, 53, 57 Advanta, 72 air-interfaced culture, 96 ALCAP see alumino calcium phosphorus oxide ceramic capsules AlgE1 enzymes, 237–8 alginases, 239 alginate-based drug delivery devices, 236–59 alginate biopolymers, 237–47 alginate oxidation reaction forming alginate dialdehyde, 245 biocompatibility, 238 calcium crosslinked alginate hydrogels, 242 chemically modified alginate, 243–5 covalently crosslinked matrices, 245–7 degradation, 239 hydrated calcium crosslinked alginate matrix, 242 interstrand crosslinking of alginate by calcium ions, 240 ionically crosslinked alginate capsules, 241

ionically crosslinked alginate hydrogels, 239–43 isomeric monomers chemical structures, 237 cells that have been immunoisolated in alginate microcapsules, 255–6 common drug delivery environments, 249 desirable attributes in the alginatebased microcapsules assembly, 253 drug delivery using alginate matrices, 247–58 composite matrices for targeted delivery, 250–1 microencapsulation for transplantation, 251–8 permeability and biointegration control, 248 use of chemically modified alginate, 249–50 future trends, 258–9 multilayered assembly applied to microcapsulate cells, 254 alginate biopolymers, 237–47 alginate dialdehyde, 244, 246 alginate-PLL microcapsules, 238 alginates, 164 Alloderm, 197 allograft, 251 allologous grafting, 150 alloplastic, 65 AlphaCor, 8–10 alumina, 2

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Index

alumino calcium phosphorus oxide ceramic capsules, 206 AMI-228/ferumoxytrol, 283 amine silane, 247 5-aminolevulinic acid, 216 aminoPVAlc-SPIONs, 225 amniotic fluid and placental stem cells, 152–3 angiogenesis, 28 induction, 317 promotion, 318–20 anti-epidermal growth factor receptor, 215 anti-protein kinase C alpha antibody, 225 Apligraft, 151 ARPE-19 cells, 257 arteries, 118–19 arteriovenous graft model, 136 articular cartilage, 57 artificial cornea, 98 artificial corneal implantation, 98–9 artificial disc replacement, 371 artificial kidney dialysers, 149 artificial nose, 60 artificial pacemakers, 149 Ascophyllum nodosum, 237 aseptic loosening, 31 asialoglycoprotein receptor, 225 ASTM D 3029–78, 336 atherosclerotic vascular diseases, 116 autograft, 350 autografting, 150 autologous coronary bypass surgery, 165 autologous fat tissue, 64 autologous tissue engineering, 150 avidin–biotin system, 275 Azobacter vinelandii, 237 balling effect, 364 balloon angioplasty, 165 basic multicellular unit, 24, 25 BD Matrigel, 258 BEL-7402, 211 bio-glasses, 159

bioabsorbables, 66 bioactive ceramics, 356 BioAir, 207 bioceramics, 206, 355 biocompatibility biomaterials and components relevant to tissue compatibility in vivo assessment, 306 drug delivery and tissue engineering, 301–21 engineered soft tissue created by stem cells, 19–39 bone: from tissue to molecular organisation, 20–2 bone development, 22–4 bone homeostasis, 24–6 bone repair after injury, 26–9 challenges of titanium implants, 30–1 current treatment options and total joint replacements, 29–30 cytokines and implants, 34–5 future trends, 37–9 growth factor delivery, 35–7 joint and bone disease, 29 titanium modifications for improved integration, 32–4 immobilised biomolecules effect on platelets adhesion and fibrinogen absorption, 305 materials and medical applications, 301–7 host reactions after medical devices implantation, 302 host response characteristics, 303 surface modifications for improving biocompatibility, 304 biocompatible, 360 biocompatible coatings, 309–11 biodegradable polymers, 236 biodegradables, 66 bioedibles, 66 biointegration, 1–12, 19–20 AlphaCor artificial corneal experience, 8–10 AlphaCor implant, 9

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Index biomaterials for dental applications, 7–8 biomaterials for orthopaedics, 1–7 titanium, 5 titanium screw, 5 functionality of tissue engineering devices, 10 future trends, 11–12 percutaneous devices, 10–11 surface modification of facial membranes, 73–8 calcium phosphate coatings, 73–5 gamma irradiation-induced grafting, 75–8 biomaterials, 1–2, 354 biocompatibility and its relevance to drug delivery and tissue engineering, 301–21 biocompatibility and medical applications, 301–7 blood–material interactions, 303–4 polymeric materials biostability and biocompatibility, 306–7 surface modifications, 304–6 tissue response to materials, 302 biointegration for dental applications, 7–8 categories bioactive, 355 bioinert, 355 bioresorbable, 355 controlled drug delivery, 307–11 biocompatible coatings, 309–11 modified polymers, 308–9 polymer co-matrix system, 309 polymers, 307–8 future trends, 320–1 materials and structural properties, 357–60 mechanical properties, 358 physical and chemical properties, 357–8 surface properties, 358–60 in vivo biological properties, 360 overview, 354–7

401

rapid prototyping in biomedical engineering, 349–90 scaffold and loaded drug/growth factor in ECM and cells integration, 315–20 angiogenesis promotion, 318 enhanced neurotrophin delivery in neural prostheses, 318 induction of angiogenesis, 317 myocardial tissue engineering, 319 vascular tissue engineering, 320 tissue engineering, 311–14 scaffold materials, 313 surface engineered biomaterials, 313–14 biomedical engineering rapid prototyping of biomaterials, 349–90 interconnected porous structure in compact and spongy bone, 375 review articles on applications of RP or SFF, 378 BioOral, 207 Bioplotter, 368 bioreactors, 98, 108, 133, 161–2, 196 bioresorbables, 66 BioSanter Pharmaceuticals, 207 biotin sandwich approach, 319 biotinylation, 319 bladder reconstruction, 167–8 blood–brain barrier, 166 BODIPY, 288 bone development, 22–4 and joint disease, 29 repair after injury, 26–9 after titanium placement, 27 from tissue to molecular organisation, 20–2 bone homeostasis, 24–6, 29 bone marrow, 20–1, 152 bone marrow stem cells, 317 bone mineral density, 329 bone morphogenetic proteins, 21, 28, 34 bone morphogenic protein-2, 317 bone remodelling, 4

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Index

bone tissue engineering, 164 borax, 250 boron neutron capture therapy, 276 Bowman’s membrane, 87 brown adipose tissue, 58 calcium alginate–polylysine–alginate microcapsule, 251 calcium phosphate, 2, 33, 159 calcium phosphate coatings, 73–5 calcium phosphate nanoparticles, 206–13 cancer chemotherapy applications, 211 gene delivery applications, 209–10 oral insulin delivery applications, 207–9 CaP nanoparticles micrograph, 209 glycemic effect of single oral dose of CAPIC, 208 tissue engineering applications, 211–13 calcium phosphate-PEG-insulin-casein, 208 canaliculi, 20 CAPIC see calcium phosphate-PEG-insulin-casein carbodiimide chemistry, 243–4 5-carboxyl-fluorescein, 220 cardiac tissue engineering see myocardial tissue engineering cartilage, 56 cartilage tissue engineering, 163–4 cationic liposomes, 217 cell sheet technology, 104–8 cell spreading, 212 ceramic coatings, 7 ceramics, 2, 159 ceramide, 211 ceruloplasmin-conjugated nanoparticles, 223 CHARITÉ Artificial Disc, 372 chitosan, 98 chloro-1-methylpyridinium chloride, 244 chlorotoxin, 220, 221

circular dichroism, 208 cisplatin, 211 citrate reduction, 213–14 coatings, 6 collagen, 64, 98, 183 collagen hydrogels, 164 Collagen IV, 305 collagen scaffolds, 125 collagen type I, 125 colloid gold see gold nanoparticles colloidal gold labelling technique, 205 Combidex, 283 commercially pure titanium implants, 7 compensatory hyperplasia, 161 computational topology design, 376 construct, 312 coralline hydroxyapatite, 251 cornea, 86 characteristics and its regeneration, 87–90 biophysical properties, 88–90 construction, 87–8 layers of cornea, 88 corneal graft rejection, 95 corneal hydration, 90 corneal regeneration, 163 corneal tissue engineering, 86–109 approaches, 95–108 active scaffold, 97–101 artificial cornea lenticules, 99 cell-seeded biodegradable scaffold-based tissue engineering, 95–7 cell sheet technology, 104–8 epithelial cell sheet preparation and transplantation procedure, 106 specimen taken on the 60th day, 99 specimen taken on the 120th day, 100 stem-cell tissue engineering, 102–4 corneal wound healing and tissue regeneration, 90–5

© Woodhead Publishing Limited, 2010

Index future trends, 108–9 human cornea characteristics and its generation, 87–90 biophysical properties, 88–90 construction, 87–8 layers of cornea, 88 corrosion, 333–4 crack initiation and propagation, 330 craniofacial skeleton, 52–3, 55, 56 Cy3 amidite, 211 cyclosporine, 149 cysteine, 243 cytokines, 26, 34–5, 38 d-mannuronic acid, 237 Dacron grafts, 117, 165, 305 daunorubicin, 217 decellularised scaffolds, 129–31 dedifferentiation, 152 defensins, 91 dermatan, 91 dermis, 58 Descemet’s membrane, 88 designer implant, 11 designer vessels, 165 dexamethasone 21-acetate, 220 dexamethasone acetate, 219 Dexon, 66 diabetic tissue engineering, 165–6 dimethyldioctadecylammonium bromide, 218 direct light fabrication, 362 direct metal laser sintering, 362 see also selective laser melting DNAzymes, 218 dodecylamine, 244 Dow Corning Silicone breast implant, 328 doxorubicin, 217, 220, 221–2, 222 drug delivery biomaterials for controlled drug delivery, 307–11 degradable polymers, 308 materials biocompatibility, 301–21 polymers, 308

403

taxol and heparin from PEG-coated PLA–chitosan co-matrix system, 310 drug targeting, 269– 82 active targeting, 274–82 commonly used protein/ligand molecules, 274 intracellular delivery and subcellular distribution, 275–7 pH-responsive nanocarrier systems, 277–8 stimuli-responsive nanocarriers, 277 temperature-responsive nanocarrier systems, 278, 282 barriers, 269–71 passive targeting, 272–3 drug accumulation in tumour site, 272 various approaches schematic representation, 271 dumbbell-like nanoparticles, 286–7 ear, 60 EC-seeding, 120 ECM see extracellular matrix egg-box model, 239, 240 elastin, 118–19, 183, 223 elastin scaffolds, 126 electrodeposition, 385 electron beam melting, 366–7, 384–5 embryonic-like stem cells, 152 embryonic stem-cell technology, 109 embryonic stem cells, 152, 153–4 see also human embryonic stem cells endochondral bone, 56 endochondral ossification, 23 endosome, 276 endosteum, 21 endothelial cells, 92 endothelial progenitor cells, 123, 135 engineered heart tissue, 163 enhanced permeability and retention effect, 272, 273, 288 EPC-seeding, 123

© Woodhead Publishing Limited, 2010

404

Index

EPCs see endothelial progenitor cells epidermal growth factor, 91 epidermal growth factor receptor, 289 epidermal growth factor receptor antibody, 286–7 epidermis, 58 epimerisation, 237 ePTFE see expanded polytetrafluoroethylene ethylenediamine, 247 EU-COST project, 252 expanded polytetrafluoroethylene, 69, 70–1, 117, 165 external elastic lamina, 119 extracellular matrix, 158–9, 197, 313, 315–20 eyes, 59, 163 facial cosmetic surgery, 63 facial membranes surface modification for optimal biointegration, 73–8 calcium phosphate coatings, 73–5 gamma irradiation-induced grafting, 75–8 facial reconstruction materials used in traditional interfacial repair, 60–73 bioresorbable and nonbiodegradable materials, 65–7 expanded polytetrafluoroethylene, 69 naturally derived materials, 64–5 polymeric materials used as tissue fillers or in soft tissue repair, 61–2 polytetrafluoroethylenes, 67–73 replacement materials for soft tissue–bone interface, 51–78 face showing soft tissue and underlying bone structure, 53 future trends, 78 organs of special senses, 58–60 skull showing principal bones, 52 tissue cross section from skin to bone, 59

tissues at bone interface, 56–8 surface modification of facial membranes, 73–8 calcium phosphate coatings, 73–5 gamma irradiation-induced grafting, 75–8 grafted ePTFE copolymer, 78 MAEP- and MOEP-grafted ePTFE, 77 Factor XII, 304 failure analysis, 337–9 knee and hip implants, 338 TMJ implant made of Co-Cr-Mo alloy, 339 failure mechanisms medical implants during prolonged used, 326–42 non-load-bearing implants, 335–7 soft tissue implants, 336–7 fatigue, 328–9 S–N curve, 330 Ferridex, 283 fibrin, 125 fibrin scaffolds, 125–6 fibrinogen, 304 fibroblast growth factors, 22 fibronectin, 10, 91, 313 fibrosis, 302 fibrous encapsulation, 302 Fickian diffusion, 181 fine-tuning Notch, 318 finite element analysis, 339, 373 floating centre of rotation, 372 fluorescence isothiocyanate, 226 fluorescent aminoPVAlc-SPIONs, 225 5-fluorouridine, 222 folic acid, 222 functional prototype development, 362 functionalised aminoPVAlc-SPIONs, 226 fused deposition modelling, 361, 364–6 scaffold technology, 379–81 3D honeycomb porous prototype, 379

© Woodhead Publishing Limited, 2010

Index 3D interconnected controlled porosity PP–TCP composite scaffolds, 381 gamma irradiation-induced grafting, 75–8 gamma secretase inhibitor, 318 gene therapy, 162 genital tissues, 168–9 gentamicin, 251 glass transition temperature, 358 glucocerebrosidase, 251 glycosaminoglycans, 90, 91, 98, 195 gold nanoparticles, 213–18, 287–9 cancer chemotherapy applications, 215–17 functionalisation using various chemical moieties, 288 gene delivery applications, 217–18 GSH-mediated payload release, 289 surface functionalisation, 286 TEM micrograph, 214 Gore-Tex, 72 grafting, 75 grafting front mechanism, 75–6 growth factor delivery, 35–7, 37–8 growth factors, 158, 315–20 HB8521 cells, 223 HeLa cells, 222 hemagglutinating virus of Japan envelopes, 225 hematopoietic stem cell, 21 Hepama-1, 222 heparin, 181, 195, 251 heparin–chitosan spheres, 309 HER2 receptor specific antibody, 275 high molecular weight kininogen, 304 HIV-LA see human immunodeficiency virus-related lipoatropy hollow implant, 34 horseradish peroxidase, 247 HSC2 cancer cells, 289 HSC3 cancer cells, 215 human aortic smooth muscle cells, 224

405

human bone marrow stem cells, 317 human corneal endothelial cell, 108 human embryonic stem cells, 39, 104, 153–4 human immunodeficiency virus-related lipoatropy, 53, 66 human microvascular endothelial cells, 193 human serum albumin, 310 human umbilical vein endothelial cells, 195, 196 hyaluronan scaffolds, 126–7 hydrogels, 181 classification, 181–5 composite bio-synthetic hydrogels, 184–5 natural hydrogels, 183 synthetic hydrogels, 184 key benefits and drawbacks, 182 hydroxyapatite, 1–4, 6, 33, 159, 250 hydroxyapatite cement, 6 Hylaform, 65 hyperbaric oxygen therapy, 11 hyperthermia, 278 ICP-OES see inductively coupled plasma-optical emission spectroscopy immune rejection, 314 immunogold, 205 implants, 34–5 induced pluripotent stem cells, 39, 156–8 inductively coupled plasma-optical emission spectroscopy, 3 inflammatory phase, 26 inorganic nanoparticles targeted drug delivery, 204–26 calcium phosphate nanoparticles, 206–13 gold nanoparticles, 213–18 iron oxide nanoparticles, 218–26 Insight, 379 insulin-like growth factor, 25 insulin-like growth factor 1, 319 Integral Biointegrated Dental Implant System, 7

© Woodhead Publishing Limited, 2010

406

Index

integrins, 91 internal elastic lamina, 118 interpenetrating polymer network, 8, 185 intramembranous ossification, 23 iodometry, 244 iron oxide nanoparticles, 218–26, 283–7 cancer therapy applications, 219–22 gene delivery applications, 223 general drug delivery and targeting, 225–6 hypothetical magnetic drug delivery system, 219 tissue engineering applications, 223–5 islet transplantation, 166 Isolagen, 65 K-562 leukaemia cells, 217 Kahalalide F analogues, 216 keratin sulphate proteoglycans, 91 keratoblasts, 93 keratoprosthesis, 8 l-guluronic acid, 237 lactoferrin, 92, 223 Lactosorb, 65 Laminaria digitata, 237 Laminaria hyperborea, 237 Laminaria japonica, 237 laminin, 91, 305, 313 laser, 361 laser additive manufacturing, 363–4 LENS and DLF processes, 363 selective laser melting or direct laser sintering, 364 selective laser sintering, 363–4 laser-engineered net shaping, 362 layer-by-layer assembly approach, 248 LENS, 363 Lewis lung carcinoma, 288 linear alipathic polyesters, 160 Link SB Charité III, 372 liquid-covered culture, 96 liver regeneration, 161–2 load-bearing implants, 327 LoVo cancer cells, 217

low critical solution temperature, 282 Lycurgus Cup, 213 lysozyme, 92 Macrocystic pyrifera, 237 MAEP see monoacryloxyethyl phosphate magnesium, 212 magnetic nanoparticles, 283–7 functionalisation, 284 functionalised nanoparticles and their applications, 285 magnetoliposomes, 225 maleimidyl 3-succinimidopropionate ligand, 220 mandril, 133 mannan, 225 Matrigel, 196 matrix proteins, 91 maturation phase, 28 MAVERICK, 371 MC 38 colon carcinoma, 288 MCF-7 breast cancer cells, 220 MC3T3-E1 preosteoblast cell, 213 medical implant failure long-term use, 326–42 clinical factors, 334–5 corrosion, 333–4 failure analysis, 337–9 failure mechanisms, 327 manufacturing deficiencies, 327–8 mechanical factors, 328–9 non-load bearing implants, 335–7 wear, 328–33 membranous bone, 56 mesenchymal stem cells, 39, 136, 152 metals, 159 methacrylic anhydride, 247 methacryloxyethyl phosphate, 76 methotrexate, 221, 284 MG-63 cells, 222 microsphere suspension fillers, 67 mineralisation phase, 28 Modelmaker II, 383 Modelworks, 383

© Woodhead Publishing Limited, 2010

Index

407

MOEP see methacryloxyethyl phosphate monoacryloxyethyl phosphate, 76 monocrystalline iron oxide nanoparticles, 224 monolayer protected clusters, 287 multifunctional nanocarrier systems, 282–9 multiple angle laser light scattering, 239 muscles, 57 MycroMesh Plus, 72 myocardial infarction, 162 myocardial tissue engineering, 162–3, 319

NanOss bone void filler, 211 nanotechnology, 205 nanoworm, 221, 224 natural polymers, 160 Nd:YAG lasers, 364 neural tissue engineering, 166–7 neuronal stem cells, 152 neuronal tissue engineering see neural tissue engineering neuropeptides, 92 New-Fill, 66 nitric oxide, 124 non-load-bearing implants, 327 nose, 59–60

N-(2-hydroxypropyl methacrylamide), 278 N-hydroxyl succinimide, 246 N-hydroxysuccinimide-folate, 226 N-isopropylacrylamide-co-acrylamide hydrogel, 215 nanobioconjugates, 205 nanobiotechnology, 12, 267 nanocarrier systems, 268 see also specific nanocarrier system multifunctional surfaces, 276 stimuli-sensitive nanocarriers for drug and gene delivery, 279–81 nanodrug delivery system, 268 NanoMarkets, 205 nanomaterials, 267 nanoparticles commonly used and their medical applications, 270 drug targeting, 269– 82 active targeting, 274–82 barriers, 269– 71 passive targeting, 272–3 folate-mediated drug targeting, 277 multifunctional nanocarrier systems, 282–9 gold nanoparticles, 287–9 magnetic nanoparticles, 283–7 systems in drug delivery, 269 targeted drug delivery, 267–89 nanospheres, 221, 224, 310

organ regeneration stem cells, 147–69 tissue engineering and stem cells, 161–9 bladder, 167–8 bone tissue engineering, 164 corneal regeneration and tissue engineering of the eye, 163 diabetic tissue engineering, 165–6 genital tissues, 168–9 myocardial tissue engineering, 162–3 neural tissue engineering, 166–7 prospects for liver regeneration, 161–2 urethra tissue engineering, 167 vascular and valvular tissue engineering, 164–5 tissue engineering basic components, 149–61 cell sources and types, 149–54 general scaffolds and materials, 158–61 growth factors, 158 reprogrammed somatic induced pluripotent stem cells, 156–8 therapeutic cloning, 154–5 organogenesis, 148 osseointegration, 3, 4–5, 6 osteoarthritis, 29 osteoblasts, 20

© Woodhead Publishing Limited, 2010

408

Index

osteoclasts, 20 osteocytes, 20 osteopontin, 6 osteoporosis, 29, 350 osteosarcomas, 29 Ostim, 212 paclitaxel, 124, 220 patient health, 334 PEG see poly(ethylene glycol) PEG spacer, 288 percutaneous devices, 10–11 periosteum, 21 peritoneal cavity, 133–5 Peyer’s patch, 207, 310 pH-responsive nanocarrier systems, 277–8 phallic reconstructive surgery, 168 phosphatidylcholine liposomes, 217 plastic surgery, 53 platelet-derived growth factor, 34, 320 platelets, 304 Pluronic F127, 220, 286 PMMA see poly(methyl methacrylate) Poloxamer 407, 273 poly(2-hydroxyethyl methacrylate), 8, 11 polyacrylic acid, 250 poly(β-amino ester), 278 polycaprolactone, 380, 383 polycaprolactonehydroxyapatite, 380 poly(carbonate urethanes), 306 poly(D,L-lactide), 381 polyε-caprolactone, 128–9 poly(ether urethanes), 306 polyethylene, 6 poly(ethylene glycol), 247, 273, 282, 309, 310, 318 polyethylene glycol-modified liposomes, 217 poly(ethylene glycol)-polycaprolactone, 318 poly(ethylene glycol)-poly(lactic acid), 318 polyethylene glycol–gold nanoparticles, 216

poly(ethylene oxide)-poly(propylene oxide)-poly(ethylene oxide), 224 poly(ethylene vinyl acetate), 314 poly(glycolic acid), 127, 160 polygycolide, 355 polyhydroxybutyrate, 355 poly(L-lactic acid), 382 poly(lactic acid), 128, 160, 309, 314 poly(lactic-co-glycolic acid), 35–6, 38, 166, 193, 317, 318 polylactide, 66, 67, 355 polymer co-matrix system, 309 polymers, 2, 307–8 poly(methyl methacrylate), 67, 160, 335–6, 357 poly(N-isopropylacrylamide), 282 poly(tetrafluoroethylene), 6, 67–73 helical structure, 68 polyurethane, 10, 306 poly(vinyl alcohol), 221 porosity polypropylene–tricalcium phosphate, 380 porous scaffolds, 352–4 conventional manufacturing techniques, 353 precision extrusion deposition, 369, 385 pressure-assisted micro-syringe, 368–9 primaquine, 250 PRODISC, 372 ProE, 379 proliferative phase, 28 Proplast I, 71 Proplast II, 71 prosthetic grafts limitations, 117 seeding, 120, 123–31 cell seeding, 120, 123–4 graft seeding with small molecules, 124 protein coating, 124 protein drug delivery facilitating materials, 179–98 incorporation of proteins/polymers into synthetic hydrogel networks, 180 protein delivery, 181

© Woodhead Publishing Limited, 2010

Index hydrogel classification, 181–5 composite bio-synthetic hydrogels, 184–5 key benefits and drawbacks, 182 natural hydrogels, 183 synthetic hydrogels, 184 protein encapsulation and release influencing factors, 185–91 decreasing mesh size, 187 network structure, 186–7 partitioning and complexation, 189–90 polymer chemistry, 188–9 polymer–protein conjugates, 190–1 tissue engineering applications, 191–7 implanted constructs vascularisation, 192–7 physiological wound healing vascularisation, 192 vascularised constructs production, 194 proteoglycans, 56, 119 Pseudomonas sp., 237 PTFE see poly(tetrafluoroethylene) quantitative computed tomography, 329 rapid prototyping (RP) electron beam melting, 366–7 fused deposition modelling, 364–6 mould design, 367 process schematics, 365 schematics of fabrication of a controlled porosity, 366 historical perspective on the development of laser-based RP, 362 laser additive manufacturing, 363–4 LENS and DLF processes, 363 selective laser melting, 364 selective laser sintering, 363–4 laser-based (see laser additive manufacturing) precision extrusion deposition, 369 net shaping RP processes, 369

409

pressure-assisted micro-syringe, 368–9 progress in tissue engineering, 386 structures of biological materials, 349–90 biomedical engineering, 373–7 designing structural implants, 369–73 material properties, 357–60 mimicking structural intricacies, 377–86 novel manufacturing approach, 360–9 overview, 354–7 patient-specific customised scaffolds, 387 three-dimensional bioplotter, 368 three-dimensional fibre deposition, 368 three-dimensional printing, 367–8 recurrent luminal narrowing, 308 regeneration, 26 repair and regeneration, 54 reproductive cloning, 155 Resovist, 283 restenosis, 308 Restylane, 65 reticuloendothelial system, 271, 272 RF magnetron sputtering, 74 rheumatoid arthritis, 29 Rhinoceros software, 383 rhinoplasty, 148 Rhodamine-labelled Pluronic/chitosan nanocapsules, 222 robotic dispensing, 384 scaffolds, 158–61, 351 corneal tissue engineering cell-seeded biodegradable, 95–7 corneal tissue regeneration in vivo, 97–101 decellularised, 129–31 allogeneic decellularised vascular, 131 xenogeneic decellularised nonvascular, 130–1

© Woodhead Publishing Limited, 2010

410

Index

xenogeneic decellularised vascular, 130 natural macromolecules for in vitro tissue engineering, 125–7 collagen scaffolds, 125 elastin scaffolds, 126 fibrin scaffolds, 125–6 hyaluronan scaffolds, 126–7 patient-specific customised via RP, 387 3D reconstructions of CT scans or histological sections, 388 resorbable ceramic mandible scaffolds, 389 synthetic macromolecules for in vitro tissue engineering, 127–9 polyε-caprolactone-based scaffolds, 128–9 polyglycolic acid-based scaffolds, 127–8 polylactic acid-based scaffolds, 128 Schiff’s reaction, 250 Sculptra, 66 selective laser melting, 362, 364 selective laser sintering, 363–4, 376, 383–4 self-assembled vascular grafts in vitro engineering, 132 in vivo engineering, 132–5 peritoneal cavity as in vivo bioreactor, 133–5 Sparks’ mandril, 133 serine proteinases, 92 Sertoli cells, 254 shape memory alloys, 357–8 silica–gold nanoshells, 215 sirolimus, 124 size exclusion chromatography, 239 skeletal device design, 369–73 new design and development, 374 orthopaedic design, 373 skin, 58 smart scaffold, 312 SMMC-7721, 222

smooth muscle cells, 196 smooth muscle progenitor cells, 135 soft-tissue–bone interface replacement materials for facial reconstruction, 51–78 facial reconstruction, 55–60 future trends, 78 materials used in traditional interfacial repair, 60–73 surface modification of facial membranes for optimal biointegration, 73–8 soft tissue implants, 336–7 SoftForm, 72 solid freeform fabrication, 354 somatic cell nuclear transfer, 154–5 somatic cells, 150–1 somatic nuclear transfer, 314 somatic stem cells, 151–2 Sparks’ mandril, 133 SPION see superparamagnetic iron oxide nanoparticles spray coating technique, 314 stem-cell tissue engineering, 102–4 stem cells, 19–39, 135–6, 169 organ regeneration, 147–69 tissue engineering and stem cells, 161–9 tissue engineering basic components, 149–61 stereolithography, 376, 381–2 PDLLA network scaffolds, 382 stimuli-responsive nanocarriers, 277 stress shielding, 31 stromal keratocytes, 92 subcutaneous fat, 58 superparamagnetic iron oxide, 277 superparamagnetic iron oxide nanoparticles, 221–2, 223 superparamagnetism, 219 surface modification, 32–3, 73–8 surface profilometry, 337 surgical errors, 334–5 synaptotagmin-I, 284 synthetic biomaterials, 30 synthetic polymers, 160–1

© Woodhead Publishing Limited, 2010

Index 6-TAMRA, 222 targeted drug delivery functionalised nanoparticles, 267–89 drug targeting, 269–82 multifunctional nanocarrier systems, 282–9 inorganic nanoparticles, 204–26 calcium phosphate nanoparticles, 206–13 gold nanoparticles, 213–18 iron nanoparticles, 218–26 taxol, 309 TCP see tricalcium phosphate tear proteins, 92 TEBV see tissue engineered blood vessels temperature-responsive nanocarrier systems, 278 tetrabutyl ammonium salt, 244 thiol conjugation, 217 thiolated alginate, 250 thiomers, 249–50 three-dimensional fibre deposition, 368, 384 three-dimensional printing, 367–8, 382–3 3D PLLA scaffolds from reverse solid-freeform fabrication, 383 thrombin, 310 thrombus, 303 thyroid cells, 149 Ti-6Al-4V, 358, 384 tissue-engineered blood vessels, 117, 118–19, 134, 137 tissue-engineered vascular grafts, 116–38, 224 tissue engineering, 10, 117, 120, 147, 351 basic components, 149–61 cell sources and types, 149–54 general scaffolds and materials, 158–61 growth factors, 158 reprogrammed somatic induced pluripotent stem cells, 156–8 somatic cell nuclear transfer, 154–5 biomaterials, 311–14

411

growth factors delivered to promote tissue formation, 316 materials biocompatibility, 301–21 scaffold and loaded drug/growth factor in ECM and cells integration, 315–20 scaffold materials, 313 and stem cells in organ regeneration, 161–9 bladder, 167–8 bone, 164 cartilage tissue, 163–4 diabetic tissue engineering, 165–6 eye and corneal regeneration, 163 genital tissues, 168–9 liver regeneration prospects, 161–2 myocardial tissue engineering, 162–3 neural tissue engineering, 166–7 urethra, 167 vascular and valvular tissue, 164–5 in vitro, 125–9 tissue fillers, 61–2 tissue necrosis factor, 216 titanium, 2, 4, 19, 38 biointegration, 34–5 biointegration showing lower roughness, 5 current modification for improved integration, 32–4 titanium implants, 7, 19, 30 current challenges, 30–1 hollow implants, 37 titanium oxide, 30 titanium screw, 5 total hip replacement, 358 total joint replacements, 29–30 trans-differentiation, 152 transforming growth factor β, 21, 28, 34 transforming growth factor-β1, 25, 26, 34 transforming growth factor-β3, 317 Trastuzumab, 275 tricalcium phosphate, 33, 381 tunica intima, 118 tunica media, 119

© Woodhead Publishing Limited, 2010

412

Index

ultra high molecular weight polyethylene, 2, 358, 371, 372 ultrasmall superparamagnetic iron oxide, 224 UltraSoft, 72 urea phosphate method, 243 urethra tissue engineering, 167 vaginal reconstruction, 169 valvular tissue engineering, 164–5 vaporisation, 364 vascular endothelial growth factor, 249, 317, 320 vascular grafts, 116–38 vascular tissue engineering, 164–5, 320 approaches, 120–36 decellularised scaffolds, 129–31 evaluated strategies, 121–2 exclusively autologous blood vessels, 131–5 natural macromolecules scaffolds, 125–7 non-absorbable prosthetic grafts seeding, 120, 123–4 stem cells, 135–6 synthetic macromolecules scaffolds, 127–9 TEBV morphology, 134 future trends, 136–8 hurdles on the way to widespread clinical use, 136–8 validation in relevant large animal models, 136 small-diameter vascular grafts, 116–38 approaches, 120–36 clinical perspective, 117 future trends, 136–8 global burden of vascular disease, 116 prosthetic grafts limitations, 117

tissue engineered blood vessels required characteristics arterial wall and its function, 118 native arteries anatomy and function, 118–19 required qualities, 119 vascularisation and protein drug delivery facilitating materials, 179–98 hydrogel classification, 181–5 incorporation of protein/polymers into synthetic polymer hydrogel networks, 180 protein delivery, 181 protein encapsulation and release, 185–91 tissue engineering applications implanted constructs, 192–7 physiological wound healing, 192 strategies for producing vascularised constructs, 194 vaso vasorum, 137 viral vectors, 209–10 viscoelasticity, 358 Vitek temporomandibular joint, 328 wear, 328–33 alumina hip implant showing stripe wear, 332 polyethylene liner from a knee implant, 331 stripe wear on ceramic hip implant, 332 Western blot method, 225 white adipose tissue, 57–8 X-ray photoelectron spectroscopy, 7 X-ray spectrographic microanalysis, 334, 337 xenograft, 251 xenografting, 150 zinc, 207, 212

© Woodhead Publishing Limited, 2010